Scanning system

ABSTRACT

An example particle therapy system includes: a particle accelerator to output a beam of charged particles; and a scanning system to scan the beam across at least part of an irradiation target. An example scanning system includes: a scanning magnet to move the beam during scanning; and a control system (i) to control the scanning magnet to produce uninterrupted movement of the beam over at least part of a depth-wise layer of the irradiation target so as to deliver doses of charged particles to the irradiation target; and (ii) to determine, in synchronism with delivery of a dose, information identifying the dose actually delivered at different positions along the depth-wise layer.

This patent application is a continuation of U.S. patent applicationSer. No. 14/184,990, entitled “SCANNING SYSTEM FOR A PARTICLE THERAPYSYSTEM”, which was filed on Feb. 20, 2014. The content of U.S. patentapplication Ser. No. 14/184,990 is hereby incorporated by reference intothis application as if set forth herein in full.

TECHNICAL FIELD

This disclosure relates generally to a particle beam scanning system foruse, e.g., with a particle therapy system.

BACKGROUND

Particle therapy systems use an accelerator to generate a particle beamfor treating afflictions, such as tumors. In operation, particles areaccelerated in orbits inside a cavity in the presence of a magneticfield, and are removed from the cavity through an extraction channel. Amagnetic field regenerator generates a magnetic field bump near theoutside of the cavity to distort the pitch and angle of some orbits sothat they precess towards, and eventually into, the extraction channel.A beam, comprised of the particles, exits the extraction channel.

A scanning system is down-beam of the extraction channel. In thiscontext, “down-beam” means closer to an irradiation target (here,relative to the extraction channel). The scanning system moves the beamacross at least part of the irradiation target to expose various partsof the irradiation target to the beam. For example, to treat a tumor,the particle beam may be “scanned” over different cross-sectional layersof the tumor.

SUMMARY

An example particle therapy system comprises: a particle accelerator tooutput a beam of charged particles; and a scanning system to scan thebeam across at least part of an irradiation target. An example scanningsystem comprises: a scanning magnet to move the beam during scanning,where a position of the beam corresponds to a current of the scanningmagnet; and a control system (i) to control the current in order toproduce uninterrupted movement of the beam across at least part of anirradiation target to deliver doses of charged particles, (ii) forpositions at which the particle beam delivers dose, to store informationidentifying a location and an amount of dose delivered, (iii) to comparea cumulative dose delivered at each position to a target cumulativedose, and (iv) if the cumulative dose does not match the targetcumulative dose at specific positions, control the current in order tomove the beam so as to deliver additional dose to the specificpositions. The example particle therapy system may include one or moreof the following features, either alone or in combination.

An example particle accelerator includes an accelerator that may beconfigured to output pulses of charged particles in accordance with aradio frequency (RF) cycle. The pulses of charged particles form thebeam. Movement of the beam across the at least part of an irradiationtarget may not be dependent upon the RF cycle. The control system may beconfigured to measure the cumulative dose delivered at each position.The measuring may be substantially synchronous with the RF cycle. Thecontrol system may be configured to measure the cumulative dosedelivered at each position. The measuring may be substantiallysynchronous with delivery of dose at each position.

The information identifying a location and an amount of dose deliveredmay comprise an amount of dose delivered at each position and at leastone of: a location of each position within the irradiation target or amagnet current corresponding to each position within the irradiationtarget. The location may correspond to three-dimensional coordinateswithin the irradiation target.

The particle therapy system may comprise: memory to store a treatmentplan that identifies, for each position, a target cumulative dose of theparticle beam. The treatment plan may omit information about individualdoses delivered to individual positions during scanning.

The scanning system may comprise: a degrader to change an energy of thebeam prior to output of the beam to the irradiation target. The degradermay be down-beam of the scanning magnet relative to the particleaccelerator. The control system may be configured to control movement ofat least part of the degrader into, or out of, a path of the beam inorder to affect the energy of the beam and thereby set a layer of theirradiation target to which charged particles are to be delivered.

The particle accelerator may comprise an ion source to provide plasmafrom which pulses in the beam are extracted. During at least part of themovement of the degrader, the ion source may be deactivated.

The particle accelerator may comprise: an ion source to provide plasmafrom which pulses in the beam are extracted; and a voltage source toprovide a radio frequency (RF) voltage to a cavity to accelerateparticles from the plasma. The cavity may have a magnetic field forcausing particles accelerated from the plasma column to move orbitallywithin the cavity. During at least part of the movement of the degrader,the voltage source may be deactivated. During the at least part of themovement of the degrader, the particle source may be deactivated at allor part of a same time that the voltage source is deactivated.

The particle accelerator may be a variable-energy particle accelerator.The control system may be configured to set an energy level of theparticle accelerator prior to scanning. The control system may beconfigured to set an energy level of the particle accelerator duringscanning.

For a position at which the particle beam delivers dose, each individualdelivery of dose may be a percentage of the total cumulative dose. Thepercentage may be less than 100% of the total cumulative dose. Thepercentage may be about, or exactly, 100% of the total cumulative dose.

The scanning magnet may have an air core, a ferromagnetic core, or acore that is a combination of air and ferromagnetic material.

Another example particle therapy system comprises: a particleaccelerator to output a beam of charged particles; and a scanning systemto scan the beam across at least part of an irradiation target. Anexample scanning system comprises: a scanning magnet to move the beamduring scanning; and a control system (i) to control the scanning magnetto produce uninterrupted movement of the beam over at least part of adepth-wise layer of the irradiation target so as to deliver doses ofcharged particles to the irradiation target; and (ii) to determine, insynchronism with delivery of a dose, information identifying the doseactually delivered at different positions along the depth-wise layer.

The example particle accelerator may be configured to output pulses ofcharged particles in accordance with a radio frequency (RF) cycle. Thepulses of charged particles form the beam. Movement of the beam may notbe dependent upon the RF cycle.

Another example particle therapy system comprises: a particleaccelerator to output a beam of charged particles; and a scanning systemto scan the beam across at least part of an irradiation target. Anexample scanning system comprises: a scanning magnet to move the beamduring scanning, where a position of the beam corresponds to a currentof the scanning magnet; and an open loop control system (i) to controlthe current to produce uninterrupted movement of the particle beamacross at least part of a layer of an irradiation target, (ii) torecord, in synchronism with delivery, doses of the particle beamdelivered to the irradiation target and at least one of: coordinates atwhich the doses were delivered or magnet currents at which the doseswere delivered, and (iii) to compensate for deficiencies in the recordeddoses relative to corresponding target cumulative doses. The exampleparticle therapy system may include one or more of the followingfeatures, either alone or in combination.

The particle accelerator may comprise: a voltage source to provide aradio frequency (RF) voltage to a cavity to accelerate particles from aplasma column, where the cavity has a magnetic field for causingparticles accelerated from the plasma column to move orbitally withinthe cavity; an extraction channel to receive the particles acceleratedfrom the plasma column and to output the received particles from thecavity towards the scanning system; and a regenerator to provide amagnetic field bump within the cavity to thereby change successiveorbits of the particles accelerated from the plasma column so that,eventually, particles output to the extraction channel. The magneticfield may be between 4 Tesla (T) and 20 T and the magnetic field bumpmay be at most 2 Tesla. The uninterrupted movement of the particle beamacross the at least part of the layer of the irradiation target may notbe dependent upon the RF frequency.

The scanning magnet may comprise an air core. The particle therapysystem may comprise a gantry on which the particle accelerator and thescanning system are mounted. The gantry may be configured to move theparticle accelerator and the scanning system around the irradiationtarget. The current of the scanning magnet may be adjusted based on aposition of the gantry.

The particle accelerator may comprise a synchrocyclotron. Uninterruptedmovement of the particle beam may occur across an entirety of the layeror across less than an entirety of the layer.

The particle therapy system may comprise a current sensor associatedwith the scanning magnet. Recording coordinates at which the doses weredelivered may comprise sampling an output of the current sensor andcorrelating the output to coordinates. The particle therapy system maycomprise an ionization chamber between the scanning magnet and theirradiation target. Recording doses of the particle beam delivered tothe irradiation target may comprise sampling an output of the ionizationchamber for each dose.

An example proton therapy system may include any of the foregoingparticle accelerators and scanning systems; and a gantry on which theparticle accelerator and scanning system are mounted. The gantry may berotatable relative to a patient position. Protons may be outputessentially directly from the particle accelerator and through thescanning system to the position of an irradiation target, such as apatient. The particle accelerator may be a synchrocyclotron.

Two or more of the features described in this disclosure, includingthose described in this summary section, may be combined to formimplementations not specifically described herein.

Control of the various systems described herein, or portions thereof,may be implemented via a computer program product that includesinstructions that are stored on one or more non-transitorymachine-readable storage media, and that are executable on one or moreprocessing devices (e.g., microprocessor(s), application-specificintegrated circuit(s), programmed logic such as field programmable gatearray(s), or the like). The systems described herein, or portionsthereof, may be implemented as an apparatus, method, or electronicsystem that may include one or more processing devices and computermemory to store executable instructions to implement control of thestated functions.

The details of one or more implementations are set forth in theaccompanying drawings and the description below. Other features,objects, and advantages will be apparent from the description anddrawings, and from the claims.

DESCRIPTION OF THE DRAWINGS

FIGS. 1 and 2 are a cross-sectional views of an example synchrocyclotronfor use in a particle therapy system.

FIG. 3 is a side view of an example scanning system.

FIG. 4 is a perspective view of components of an example scanningsystem.

FIG. 5 is a front view of an example magnet for use in a scanning systemof the type shown in FIGS. 3 and 4.

FIG. 6 is a perspective view of an example magnet for use in a scanningsystem of the type shown in FIGS. 3 and 4.

FIG. 7 is a perspective view of an example energy degrader (rangemodulator) for use in a scanning system of the type shown in FIGS. 3 and4.

FIG. 8 is a perspective view of a process for moving a plate of anenergy degrader in the path of a particle beam.

FIG. 9 is a flowchart showing an example process for performing rasterscanning that may be performed using the hardware of FIGS. 1 to 8.

FIG. 10 is a top view show an example cross-section of an irradiationtarget and a radiation scan path.

FIG. 11 is a perspective view of an example therapy system.

FIG. 12 is an exploded perspective view of components of an examplesynchrocyclotron for use in the particle therapy system.

FIG. 13 is a cross-sectional view of the example synchrocyclotron.

FIG. 14 is a perspective view of the example synchrocyclotron.

FIG. 15 is a cross-sectional view of an example ion source for use inthe synchrocyclotron.

FIG. 16 is a perspective view of an example dee plate and an exampledummy dee for use in the synchrocyclotron.

FIG. 17 shows a patient positioned within an example inner gantry of theexample particle therapy system in a treatment room.

FIG. 18 is a conceptual view of an example particle therapy system thatmay use a variable-energy particle accelerator.

FIG. 19 is a perspective, exploded view of an example magnet system thatmay be used in a variable-energy particle accelerator.

FIG. 20 is an example graph showing energy and current for variations inmagnetic field and distance in a particle accelerator.

FIG. 21 is a side view of an example structure for sweeping voltage on adee plate over a frequency range for each energy level of a particlebeam, and for varying the frequency range when the particle beam energyis varied.

Like reference symbols in the various drawings indicate like elements.

DETAILED DESCRIPTION

Described herein is an example of a particle accelerator for use in asystem, such as a proton or ion therapy system. The example particletherapy system includes a particle accelerator—in this example, asynchrocyclotron—mounted on a gantry. The gantry enables the acceleratorto be rotated around a patient position, as explained in more detailbelow. In some implementations, the gantry is steel and has two legsmounted for rotation on two respective bearings that lie on oppositesides of a patient. The particle accelerator is supported by a steeltruss that is long enough to span a treatment area in which the patientlies and that is attached at both ends to the rotating legs of thegantry. As a result of rotation of the gantry around the patient, theparticle accelerator also rotates.

In an example implementation, the particle accelerator (e.g., thesynchrocyclotron) includes a cryostat that holds one or moresuperconducting coils, each for conducting a current that generates amagnetic field (B). In this example, the cryostat uses liquid helium(He) to maintain each coil at superconducting temperatures, e.g., 4°Kelvin (K). Magnetic yokes or smaller magnetic pole pieces are locatedinside the cryostat, and define a cavity in which particles areaccelerated.

In this example implementation, the particle accelerator includes aparticle source (e.g., a Penning Ion Gauge—PIG source) to provide aplasma column to the cavity. Hydrogen gas is ionized to produce theplasma column. A voltage source provides a radio frequency (RF) voltageto the cavity to accelerate pulses of particles from the plasma column.

As noted, in an example, the particle accelerator is a synchrocyclotron.Accordingly, the RF voltage is swept across a range of frequencies toaccount for relativistic effects on the particles (e.g., increasingparticle mass) when accelerating particles from the plasma column. Themagnetic field produced by running current through a superconductingcoil causes particles accelerated from the plasma column to accelerateorbitally within the cavity. In other implementations, a particleaccelerator other than a synchrocyclotron may be used. For example, acyclotron, a synchrotron, a linear accelerator, and so forth may besubstituted for the synchrocyclotron described herein.

In the example synchrocyclotron, a magnetic field regenerator(“regenerator”) is positioned near the outside of the cavity (e.g., atan interior edge thereof) to adjust the existing magnetic field insidethe cavity to thereby change locations (e.g., the pitch and angle) ofsuccessive orbits of the particles accelerated from the plasma column sothat, eventually, the particles output to an extraction channel thatpasses through the cryostat. The regenerator may increase the magneticfield at a point in the cavity (e.g., it may produce a magnetic field“bump” at an area of the cavity), thereby causing each successive orbitof particles at that point to precess outwardly toward the entry pointof the extraction channel until it reaches the extraction channel. Theextraction channel receives particles accelerated from the plasma columnand outputs the received particles from the cavity as a particle beam.

The superconducting (“main”) coils can produce relatively high magneticfields. The magnetic field generated by a main coil may be within arange of 4 T to 20 T or more. For example, a main coil may be used togenerate magnetic fields at, or that exceed, one or more of thefollowing magnitudes: 4.0 T, 4.1 T, 4.2 T, 4.3 T, 4.4 T, 4.5 T, 4.6 T,4.7 T, 4.8 T, 4.9 T, 5.0 T, 5.1 T, 5.2 T, 5.3 T, 5.4 T, 5.5 T, 5.6 T,5.7 T, 5.8 T, 5.9 T, 6.0 T, 6.1 T, 6.2 T, 6.3 T, 6.4 T, 6.5 T, 6.6 T,6.7 T, 6.8 T, 6.9 T, 7.0 T, 7.1 T, 7.2 T, 7.3 T, 7.4 T, 7.5 T, 7.6 T,7.7 T, 7.8 T, 7.9 T, 8.0 T, 8.1 T, 8.2 T, 8.3 T, 8.4 T, 8.5 T, 8.6 T,8.7 T, 8.8 T, 8.9 T, 9.0 T, 9.1 T, 9.2 T, 9.3 T, 9.4 T, 9.5 T, 9.6 T,9.7 T, 9.8 T, 9.9 T, 10.0 T, 10.1 T, 10.2 T, 10.3 T, 10.4 T, 10.5 T,10.6 T, 10.7 T, 10.8 T, 10.9 T, 11.0 T, 11.1 T, 11.2 T, 11.3 T, 11.4 T,11.5 T, 11.6 T, 11.7 T, 11.8 T, 11.9 T, 12.0 T, 12.1 T, 12.2 T, 12.3 T,12.4 T, 12.5 T, 12.6 T, 12.7 T, 12.8 T, 12.9 T, 13.0 T, 13.1 T, 13.2 T,13.3 T, 13.4 T, 13.5 T, 13.6 T, 13.7 T, 13.8 T, 13.9 T, 14.0 T, 14.1 T,14.2 T, 14.3 T, 14.4 T, 14.5 T, 14.6 T, 14.7 T, 14.8 T, 14.9 T, 15.0 T,15.1 T, 15.2 T, 15.3 T, 15.4 T, 15.5 T, 15.6 T, 15.7 T, 15.8 T, 15.9 T,16.0 T, 16.1 T, 16.2 T, 16.3 T, 16.4 T, 16.5 T, 16.6 T, 16.7 T, 16.8 T,16.9 T, 17.0 T, 17.1 T, 17.2 T, 17.3 T, 17.4 T, 17.5 T, 17.6 T, 17.7 T,17.8 T, 17.9 T, 18.0 T, 18.1 T, 18.2 T, 18.3 T, 18.4 T, 18.5 T, 18.6 T,18.7 T, 18.8 T, 18.9 T, 19.0 T, 19.1 T, 19.2 T, 19.3 T, 19.4 T, 19.5 T,19.6 T, 19.7 T, 19.8 T, 19.9 T, 20.0 T, 20.1 T, 20.2 T, 20.3 T, 20.4 T,20.5 T, 20.6 T, 20.7 T, 20.8 T, 20.9 T, or more. Furthermore, a maincoil may be used to generate magnetic fields that are within the rangeof 4 T to 20 T (or more, or less) that are not specifically listedabove.

In some implementations, such as the implementation shown in FIGS. 1 and2, large ferromagnetic magnetic yokes act as a return for stray magneticfield produced by the superconducting coils. For example, in someimplementations, the superconducting magnet can generate a relativelyhigh magnetic field of, e.g., 4 T or more, resulting in considerablestray magnetic fields. In some systems, such as that shown in FIGS. 1and 2, the relatively large ferromagnetic return yoke 100 is used as areturn for the magnetic field generated by superconducting coils. Amagnetic shield surrounds the yoke. The return yoke and the shieldtogether dissipated stray magnetic field, thereby reducing thepossibility that stray magnetic fields will adversely affect theoperation of the accelerator.

In some implementations, the return yoke and shield may be replaced by,or augmented by, an active return system. An example active returnsystem includes one or more active return coils that conduct current ina direction opposite to current through the main superconducting coils.In some example implementations, there is an active return coil for eachsuperconducting coil, e.g., two active return coils—one for eachsuperconducting coil (referred to as a “main” coil). Each active returncoil may also be a superconducting coil that surrounds the outside of acorresponding main superconducting coil.

Current passes through the active return coils in a direction that isopposite to the direction of current passing through the main coils. Thecurrent passing through the active return coils thus generates amagnetic field that is opposite in polarity to the magnetic fieldgenerated by the main coils. As a result, the magnetic field generatedby an active return coil is able to dissipate at least some of therelatively strong stray magnetic field resulting from the correspondingmain coil. In some implementations, each active return may be used togenerate a magnetic field of between 2.5 T and 12 T or more. An exampleof an active return system that may be used is described in U.S. patentapplication Ser. No. 13/907,601, filed on May 31, 2013, the contents ofwhich are incorporated herein by reference.

Referring to FIG. 3, at the output of extraction channel 102 of particleaccelerator 105 (which may have the configuration shown in FIGS. 1 and2), is an example scanning system 106 that may be used to scan theparticle beam across at least part of an irradiation target. FIG. 4 alsoshows examples of components of the scanning system. These include, butare not limited to, a scanning magnet 108, an ion chamber 109, and anenergy degrader 110. Other components that may be incorporated into thescanning system are not shown in FIG. 4, including, e.g., one or morescatterers for changing beam spot size.

In an example operation, scanning magnet 108 is controllable in twodimensions (e.g., Cartesian XY dimensions) to direct the particle beamacross a part (e.g., a cross-section) of an irradiation target. Ionchamber 109 detects the dosage of the beam and feeds-back thatinformation to a control system to adjust beam movement. Energy degrader110 is controllable to move material (e.g., one or more individualplates) into, and out of, the path of the particle beam to change theenergy of the particle beam and therefore the depth to which theparticle beam will penetrate the irradiation target. In this way, theenergy degrader selects a depth-wise layer of an irradiation target toscan in two dimensions.

FIGS. 5 and 6 show views of an example scanning magnet 108. In thisexample implementation, scanning magnet 108 includes two coils 111,which control particle beam movement in the X direction, and two coils112, which control particle beam movement in the Y direction. Control isachieved, in some implementations, by varying current through one orboth sets of coils to thereby vary the magnetic field(s) producedthereby. By varying the magnetic field(s) appropriately, the particlebeam can be moved in the X and/or Y direction across the irradiationtarget. In some implementations, the scanning magnet is not movablephysically relative to the particle accelerator. In otherimplementations, the scanning magnet may be movable relative to theparticle accelerator (e.g., in addition to the movement provided by thegantry). In some implementations, the scanning magnet may becontrollable to move the particle beam continuously so that there isuninterrupted motion of the particle beam over at least part of, andpossibly all of, a layer of an irradiation target being scanned. Inother implementations, the scanning magnets are controllable atintervals or specific times. In some implementations, there may bedifferent scanning magnets to control all or part movement of a particlebeam in the X and/or Y directions.

In some implementations, scanning magnet 108 may have an air core. Inother implementations, scanning magnet 108 may have a ferromagnetic(e.g., an iron) core. In general, a magnet having an air core includes amagnetic coil around a core that is a non-ferromagnetic material, suchas air. For example, an air core magnet may include self-supportingcoils that surround air. In some implementations, an air core magnet mayinclude coils that are wound around an insulator, such as ceramic orplastic, which may or may not include air.

In some cases, an air core may have advantages over a ferromagneticcore. For example, the amount that the particle beam moves (e.g., isdeflected) in the X and/or Y directions is determined, at least in part,based on the amount of current applied to the magnet (referred to as the“magnet current”). A scanning magnet typically has a movement (ordeflection) range, which is the extent over which the magnet will movethe beam. At extremes of this range, such as at the edges, largeramounts of current are applied to the scanning magnet in order toachieve relatively high amounts of beam deflection. Some types ofscanning magnets having a ferromagnetic core may saturate at theseextremes, resulting in a non-linear relationship between current andmagnet movement. That is, the amount of deflection produced by themagnet may not be linearly proportional to the amount of current appliedto the magnet. Due to this non-linearity, in some cases, it may bedifficult to determine and/or set some beam locations using magnetcurrent. Accordingly, when a scanning magnet having a ferromagnetic coreis used, there may need to be some calibration and/or compensationperformed in order to correct for non-linearities such as that describedabove.

In contrast, a scanning magnet having an air core may not saturate inthe same manner as a scanning magnet having a ferromagnetic core. Forexample, an air core magnet may not saturate or may saturate less than amagnet having a ferromagnetic core. As a result, the relationshipbetween current and magnet movement may be more linear, particularly atthe range extremes, making determinations of beam location based onmagnet current more accurate, at least in some cases. This increasedlinearity also can enable more accurate movement of the beam,particularly at range extremes. That is, since the relationship betweencurrent and beam movement is generally more linear over a larger rangewhen an air core scanning magnet is used, beam movement may be moreeasily reproducible using an air core scanning magnet. This can beadvantageous, since a depth-wise layer of an irradiation target mayrequire multiple scans, each providing a percentage of a totalcumulative radiation dose. Precision in delivery of multiple doses tothe same area, such as that which can be obtained through use of an aircore scanning magnet, can affect the efficacy of the treatment.

Although the relationship between current and magnet movement may bemore linear in an air core magnet, in some cases, an air core magnet maybe more susceptible to stray magnetic fields than a magnet having aferromagnetic core. These stray magnetic fields may impact the scanningmagnet during motion of the scanning magnet produced by the gantry.Accordingly, in some implementations that use a scanning magnet havingan air core, the current applied to the scanning magnet to move the beammay be calibrated to account for the position of the scanning magnetrelative to the patient (or, correspondingly, to account for theposition of the gantry, since the position of the gantry corresponds tothe position of the scanning magnet relative to the patient). Forexample, the behavior of the scanning magnet may be determined and, ifnecessary, corrected, for different rotational positions (angles) of thegantry, e.g., by increasing or decreasing some applied current based onrotational position.

In some implementations, the scanning magnet may have a core that iscomprised of both air and a ferromagnetic material (e.g., iron). In suchimplementations, the amount and configuration of air and ferromagneticmaterial in the core may be determined taking the foregoing factors intoaccount.

In some implementations, a current sensor 118 may be connected to, or beotherwise associated with, scanning magnet 108. For example, the currentsensor may be in communication with, but not connected to, the scanningmagnet. In some implementations, the current sensor samples currentapplied to the magnet, which may include current to coil(s) forcontrolling beam scanning in the X direction and/or current to coil(s)for controlling beam scanning in the Y direction. The current sensor maysample current through the magnet at times that correspond to theoccurrence of pulses in the particle beam or at a rate that exceeds therate that the pulses occur in the particle beam. In the latter case, thesamples, which identify the magnet current, are correlated to detectionof the pulses by the ion chamber described below. For example, the timesat which pulses are detected using the ion chamber (described below) maybe correlated in time to samples from the current sensor, therebyidentifying the current in the magnet coil(s) at the times of thepulses. Using the magnet current, it thus may be possible to determinethe location on the irradiation target (e.g., on a depth-wise layer ofthe irradiation target) where each pulse, and thus dose of particles,was delivered. The location of the depth-wise layer may be determinedbased on the position of the energy degrader (e.g., the number ofplates) in the beam path.

During operation, the magnitude(s) (e.g., value(s)) of the magnetcurrent(s) may be stored for each location at which a dose is delivered,along with the amount (e.g., intensity) of the dose. A computer system,which may be either on the accelerator or remote from the acceleratorand which may include memory and one or more processing devices, maycorrelate the magnet current to coordinates within the radiation target,and those coordinates may be stored along with the amount of the dose.For example, the location may be identified by depth-wise layer numberand Cartesian XY coordinates or by Cartesian XYZ coordinates (with thelayer corresponding to the Z coordinate). In some implementations, boththe magnitude of the magnet current and the coordinate locations may bestored along with the dose at each location. This information may bestored in memory either on, or remote from, the accelerator. Asdescribed in more detail below, this information may be used duringscanning to apply multiple doses to the same locations to achieve targetcumulative doses.

In some implementations, ion chamber 109 detects dosage (e.g., one ormore individual doses) applied by the particle beam to positions on anirradiation target by detecting the numbers of ion pairs created withina gas caused by incident radiation. The numbers of ion pairs correspondto the dose provided by the particle beam. That information is fed-backto the computer system and stored in memory along with the time that thedose is provided. This information may be correlated to, and stored inassociation with, the location at which the dose was provided and/or themagnitude of the magnet current at that time, as described above.

As described in more detail below, in some implementations, the scanningsystem is run open loop, in which case the particle beam is moved freelyand uninterrupted across an irradiation target so as to substantiallycover the target with radiation. As the radiation is delivered,dosimetry implemented by the particle therapy control system records(e.g., stores) the amount of the radiation per location and informationcorresponding to the location at which the radiation was delivered. Thelocation at which the radiation was delivered may be recorded ascoordinates or as one or more magnet current values, and the amount ofthe radiation that was delivered may be recorded as dosage in grays.Because the system is run open loop, the delivery of the radiation isnot synchronized to the operation of the particle accelerator (e.g., toits RF cycle). However, the dosimetry may be synchronized to theoperation of the particle accelerator. More specifically, the dosimetryrecords the amount and location of each dose delivered as the dose isdelivered (that is, as close in time to delivery as possible given thelimits of technology). Since the dose is delivered in synchronism withthe operation of the accelerator (e.g., one pulse is delivered per RFcycle), in some implementations, the dosimetry that records the dose andthe location operates in synchronism, or substantially in synchronism,with delivery of radiation doses to the target, and thus in synchronismwith the operation of the particle accelerator, such as its RF cycle.

FIG. 7 shows a range modulator 115, which is an example implementationof energy degrader 110. In some implementations, such as that shown inFIG. 7, the range modulator includes a series of plates 116. The platesmay be made of one or more of the following example materials: carbon,beryllium or other material of low atomic number. Other materials,however, may be used in place of, or in addition to, these examplematerials.

One or more of the plates is movable into, or out of, the beam path tothereby affect the energy of the particle beam and, thus, the depth ofpenetration of the particle beam within the irradiation target. Forexample, the more plates that are moved into the path of the particlebeam, the more energy that will be absorbed by the plates, and the lessenergy the particle beam will have. Conversely, the fewer plates thatare moved into the path of the particle beam, the less energy that willbe absorbed by the plates, and the more energy the particle beam willhave. Higher energy particle beams typically penetrate deeper into theirradiation target than do lower energy particle beams. In this context,“higher” and “lower” are meant as relative terms, and do not have anyspecific numeric connotations.

Plates are moved physically into, and out of, the path of the particlebeam. For example, as shown in FIG. 8, a plate 116 a moves along thedirection of arrow 117 between positions in the path of the particlebeam and outside the path of the particle beam. The plates arecomputer-controlled. Generally, the number of plates that are moved intothe path of the particle beam corresponds to the depth at which scanningof an irradiation target is to take place. For example, the irradiationtarget can be divided into cross-sections or depth-wise layers, each ofwhich corresponds to an irradiation depth. One or more plates of therange modulator can be moved into, or out of, the beam path to theirradiation target in order to achieve the appropriate energy toirradiate each of these cross-sections or depth-wise layers of theirradiation target. The range modulator may be stationary relative tothe particle beam during scanning of a part of (e.g., cross-section of)an irradiation target, except for its plates moving in and out of thepath of the particle beam. Alternatively, the range modulator of FIGS. 7and 8 may be replaced with a range modulator that, at least some of thetime, tracks movement of the particle beam, thereby enabling use ofsmaller plates.

In implementations that use a range modulator of the type describedabove, the number of plates that are moved into the beam pathdetermine/set the depth-wise layer of the irradiation target that is tobe scanned. For example, if two plates are moved into the beam path, thelayer will be more shallow than if one or no plates are moved into thebeam path. The layer may be identified, and stored in memory, based onthe number of plates moved into the beam path. In some implementations,the plates may have different thicknesses. In such implementations, thethicknesses of the various plates also affect which layer is to bescanned (e.g., how deep the particle beam will penetrate the target).

In some implementations, the particle accelerator may be avariable-energy particle accelerator, such as the example particleaccelerator described in U.S. patent application Ser. No. 13/916,401,filed on Jun. 12, 2013, the contents of which are incorporated herein byreference. In example systems where a variable-energy particleaccelerator is used, there may be less need for an energy degrader ofthe type described herein, as the energy level of the particle beam maybe controlled by the particle accelerator. For example, in some systemsthat employ a variable-energy particle accelerator, an energy degradermay not be needed. In some systems that employ a variable-energyparticle accelerator, an energy degrader may still be used to changebeam energy levels.

In some implementations, a treatment plan is established prior totreating the irradiation target. The treatment plan may be stored inmemory that is accessible to a computer system that controls operationof the particle therapy system. The treatment plan may includeinformation about how radiation treatment is to be provided by theparticle therapy system. For example, the treatment plan may specify howscanning is to be performed for a particular irradiation target. In someimplementations, the treatment plan specifies that raster scanning is tobe performed. Raster scanning includes producing an uninterruptedmovement of the particle beam across the irradiation target. Forexample, the scanning magnet moves continually to scan (e.g., move) theparticle beam across the irradiation target so as to produceuninterrupted movement of the particle beam over at least part of alayer of an irradiation target. The movement may be uninterrupted acrossan entire layer of the irradiation target or across only part of alayer. In some implementations, the beam may be moved at a constantspeed along all or part of a layer of the irradiation target. In someimplementations, the speed at which the beam is moved along all or partof a layer of the irradiation target may vary. For example, the particlebeam may move more quickly across internal portions of a layer than atedges of the layer. The speed of movement may be specified in thetreatment plan.

In some implementations, the treatment plan may also specify the targetcumulative dose of radiation (particles) to be applied to variouspositions on layers of an irradiation target. The dose is cumulative inthe sense that it may be achieved through application of one or moredoses of particles. For example, the same location (e.g., in XYZ space)on an irradiation target may be irradiated ten times, each time with 10%of the target cumulative dose to achieve the target cumulative dose. Insome implementations, the treatment plan need not specify the amount ofdose for each location, the locations, or the number of times thatlocations are to be irradiated. That is, this information may be omittedfrom the treatment plan in some implementations. Rather, in someimplementations, the intensity of the particle beam may be setbeforehand to provide a particular dose of radiation per instance ofirradiation. The particle beam may then be scanned over a layer of theirradiation target in an open loop manner, without requiring feedback tomove to a next location. As the particle beam is scanned, the locationof the beam is determined and the corresponding dose at that location isdetermined. The determination may be made at about the same time as thescanning and delivery (that is, as close in time to delivery as possiblegiven the limits of technology). The cumulative dose at that location,which includes the current dose as well as any dose previously deliveredduring the current treatment, is compared to the target cumulative dosefrom the treatment plan. If the two do not match, then additional dosemay be applied to that location during a subsequent scan. Since it isnot always known precisely how much radiation will be delivered to alocation per scan, the number of times that a location is scanned maynot be set beforehand. Likewise, since there may be fluctuations in theamount of radiation actually delivered per scan to a location, theprecise amount of radiation per scan is not necessarily set beforehand.Consequently, in some implementations, such information need not beincluded in the treatment plan.

In some implementations, the treatment plan may also include one or morepatterns, over which the particle beam may be scanned per layer. Thetreatment plan may also specify the number of plates of an energydegrader to achieve a particular energy level/layer. Otherimplementations may include information in addition to, or instead of,that specified above.

In some implementations, the overall treatment plan of an irradiationtarget may include different treatment plans for differentcross-sections (layers) of the irradiation target. The treatment plansfor different cross-sections may contain the same information ordifferent information, such as that provided above.

In some implementations, the scanning system may include a collimator120 (FIG. 3) to collimate the particle bean, which may include anaperture that is placeable relative to the irradiation target to limitthe extent of the particle beam and thereby alter the shape of the spotapplied to the irradiation target. For example, the collimator may beplaced in the beam path down-beam of the energy degrader and before theparticle beam hits the irradiation target. The collimator may contain anarea (e.g., a hole or a transmissive material) through which theparticle beam passes and another material (e.g., brass) around the holethat inhibits or prevents passage of the particle beam.

In some implementations, the collimator may include a structure definingan edge. The structure may include a material, such as brass, thatinhibits transmission of the particle beam. The structure may becontrollable to move in two dimensions relative to the irradiationtarget so that at least part of the structure is between at least partof the particle beam and the irradiation target. For example, thestructure may be movable in the X and Y directions of a plane thatintersects the particle beam and that is parallel to, or substantiallyparallel to, a cross-section of the irradiation target that is beingtreated. Use of a collimator in this manner may be beneficial in that itcan be used to customize the cross-sectional shape of the particle beamthat reaches the patient, thereby limiting the amount of particle beamthat extends beyond the radiation target. Examples of collimators andenergy degraders that may be used are described in U.S. patentapplication Ser. No. 14/137,854, which was filed on Dec. 20, 2013, andwhich is incorporated herein by reference.

As noted above, in some implementations, scanning is performed in anopen-loop manner, e.g., by an open-loop control system that may beimplemented using one or more processing devices, such as the computingdevice that controls the particle therapy system. In this example,open-loop scanning includes moving the particle beam across anirradiation target to substantially cover the target with radiation. Insome implementations, movement is not synchronized with operation of theaccelerator, e.g., with the RF frequency, but rather runs independentlyof the operation of the accelerator when the accelerator is operating.For example, movement of the particle beam may be uninterrupted, and notdependent upon the RF cycle of the particle accelerator. Uninterruptedmovement may occur across all or part of a layer of an irradiationtarget. However, as described herein, the dosimetry may be synchronizedwith delivery of pulses of the particle beam to the irradiation target.In examples where the dosimetry is synchronized with delivery of pulsesof the particle beam, the dosimetry is also synchronized with operationof the accelerator (e.g., with the RF frequency used to draw pulses ofthe particle beam from the ion source plasma column).

The radiation level of an individual dose of particle beam (e.g., anindividual pulse from the accelerator) may be set beforehand. Forexample, each individual dose may be specified in grays. An individualdose may be, or correspond to, a percentage of the target cumulativedose that is to be applied to a location (e.g., an XYZ coordinate) in anirradiation target. In some implementations, the individual dose may be100% of the target cumulative dose and, as a result, only a single scanmay be needed to deliver a single dose of radiation (e.g., one or moreparticle pulses) per location to the irradiation target. In someimplementations, the individual dose may be less than 100% of the targetcumulative dose, resulting in the need for multiple scans of the samelocation to deliver multiple doses of radiation to the irradiationtarget. The individual dose may be any appropriate percentage of thetarget cumulative dose, such as: 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9%,10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, 20%, 25%, 30%, 35%,40%, 45%, 50%, 55%, 60%, 65%, 70%, 75%, 80%, 85%, 90%, 95%, or anypercentage in between these values.

The scanning magnet current may be controlled, in accordance with thetreatment plan, to scan a depth-wise layer of the irradiation target.The layer is selected by appropriately positioning one or more energydegraders from the range compensator in the path of the particle beamand/or by setting an energy level of a variable-energy particleaccelerator. As the layer is scanned, the current sensor samples thecurrent applied to the scanning magnet. The amount of magnet current maybe recorded, e.g., stored in memory. If more than one magnet or magnetcoil is used, the amount of magnet current may be stored along with theidentity of the magnet or coil. In addition, the current may becorrelated to coordinates within the irradiation target (e.g., CartesianXYZ coordinates) and those coordinates may be stored in addition to, orinstead of, the corresponding magnet current. As explained above, thecurrent sensor may sample the magnet current and correlate the samplingtime to the time at which an irradiation dose (e.g., pulse) isdelivered.

In this regard, ion chamber 109 may detect the intensity of dosesdelivered to the irradiation target as that dose is delivered. Theintensity of each dose is recorded (e.g., stored in memory) along withthe location of each delivered dose. As noted, the location of eachdelivered dose may be stored by coordinates, magnet current, or usingsome other appropriate metric. As noted above, the dosimetry—the doseverification—may be synchronized with delivery of the dose and, thus,with the output of the accelerator (which corresponds to the RFfrequency, as described above). Accordingly, in some implementations,each time a dose is delivered, the intensity of that dose is determinedalmost immediately and the location at which the dose is applied isdetermined almost immediately. This information may be stored in one ormore tables (e.g., one table per layer or multiple tables per layer) orother appropriate computer storage construct.

In some implementations, the tables may be updated as additional dosesare delivered. For example, a table may keep a running track of theamount of dose delivered at each location. So, if the beam dose is “X”grays, at a first scan pass, the table may record X grays for alocation. At a second scan pass, the table may record 2X grays, and soforth until the target cumulative dose is reached.

In this regard, for each location, a processing device associated withthe accelerator (e.g., the computer system that controls the particletherapy system) may compare the cumulative dose from a table, such asthat described above, to the target cumulative dose. If the cumulativedose matches the target cumulative dose, treatment for that location (orlayer) is deemed completed. If the cumulative dose does not match thetarget cumulative dose, additional treatment is performed. For example,the layer or location is scanned again at the same locations, which areobtained from the table. The linear correlation between magnet currentand beam movement produced by use of an air core magnet can facilitaterepeated, and relatively accurate, repeated scanning at the samelocations during multiple passes of the beam during scanning.

Scanning may be repeated, at the same locations, any appropriate numberof times until the target cumulative dose is reached at each location.In this regard, the entire layer may be re-scanned or only selectportions of the layer may be re-scanned, dependent upon the targetcumulative doses for the different locations on the layer. In someimplementations, the intensity of the particle beam is not variedbetween scans. In other implementations, the intensity of the particlebeam may be varied between scans, particularly if a small dose isrequired to top-off a cumulative dose to reach the target cumulativedose. The intensity of the dose may be increased or decreased, e.g., byaltering the operation of the ion source (e.g., increasing the plasmaionization), altering the sweep of the RF frequency, or by any otherappropriate methods. Examples of ways to vary the intensity of the doseare described in U.S. patent application Ser. No. 14/039,307, which wasfiled on Sep. 27, 2013, and which is incorporated herein by reference.

As noted, scanning may be repeated for an entire layer or for only aportion of a layer. In some implementations, an entire layer, or aportion thereof, may be fully treated before treating another layer.That is, scanning may be repeated until the total cumulative dose isreached for each location on a layer before another layer is treated. Insome implementations, each layer may be treated partially (e.g., scannedonce) in sequence, and then re-scanned in sequence. In someimplementations, several designated layers may be completely treatedbefore other layers are treated. In some implementations, the entiretarget may be scanned once, followed by successive scans of the entiretarget until the appropriate total cumulative dose is delivered to eachlocation.

During movement between layers, the beam may be turned-off. For example,during movement between layers, the ion source may be turned-off,thereby disrupting the output of the beam. During movement betweenlayers, the RF sweep in the particle accelerator may be turned-off,thereby disrupting the extraction (and thus output) of the beam. Duringmovement between layers, both the ion source and the circuitry thatcreates the RF sweep may be deactivated in some implementations. In someimplementations, rather than turning-off the ion source and/or the RFsweep during movement between layers, the beam may be deflected to abeam-absorbing material using a kicker magnet (not shown) or thescanning magnet.

Different cross-sections of the irradiation target may be scannedaccording to different treatment plans. As described above, an energydegrader is used to control the scanning depth. In some implementations,the particle beam may be interrupted or redirected during configurationof the energy degrader. In other implementations, this need not be thecase.

Described herein are examples of treating cross-sections of anirradiation target. These may be cross-sections that are roughlyperpendicular to the direction of the particle beam. However, theconcepts described herein are equally applicable to treating otherportions of an irradiation target that are not cross-sectionsperpendicular to the direction of the particle beam. For example, anirradiation target may be segmented into spherical, cubical or othershaped volumes, and those volumes may be treated using the exampleprocesses, systems, and/or devices described herein.

FIG. 9 is a flowchart showing an example implementation of the scanningprocess described herein. Although process 200 of FIG. 9 is described inthe context of the hardware described herein, process 200 may beperformed using any appropriate hardware. The operations shown inprocess 200 may be performed in the same order as depicted or in adifferent order, where appropriate.

According to process 200, a treatment plan is stored (201). Thetreatment plan may be a treatment plan as described above. For example,the treatment plan may specify the type of scanning (e.g., uninterruptedraster scanning) and the total cumulative dose of radiation to bedelivered to each location in each layer of an irradiation target. Thetreatment plan may omit, e.g., the doses to be delivered for each scanat individual locations and their intensities, as well as the number ofdoses to be delivered to each location and the identity of thelocations.

An energy degrader may be set to select (202) a layer, and current maybe applied to the magnet and controlled to move (203) the particle beamin accordance with a pattern set forth, e.g., in the treatment plan, toscan the layer. The current control may produce uninterrupted movementof the beam across at least part of the irradiation target to deliverdoses of charged particles. An example of a pattern of beam movement 230across a layer 233 of an irradiation target is shown in FIG. 10. As thebeam moves, each pulse of the beam delivers a dose of radiation to thetarget. The dose has an intensity, which may be set beforehand in theaccelerator or during scanning, and is delivered to specific positions.The exact positions at which dose is to be delivered need not be setbeforehand, but rather may be arrived at by a combination of beammovement and pulse output.

For positions at which the dose is delivered, information is stored(204) (or otherwise recorded), which identifies a location and an amountof dose delivered at the location. This information is typically storedafter the dose is delivered. As explained above, the information may bedetermined as close to delivery of the dose as possible, using the ionchamber to determine particle beam intensity (e.g., the amount of thedose) and the current sensor on the scanning magnet to determine thelocation at which the dose is delivered. As described above, in someimplementations, in synchronism with the delivery, informationidentifying doses of the particle beam delivered to the irradiationtarget is stored along with at least one of: coordinates at which thedoses were delivered or magnet currents at which the doses weredelivered. As also described above, this information may be stored intables, which may be used to store the cumulative dose of radiationapplied at positions on various layers of an irradiation target.

The entire layer may be scanned and information therefor recorded, asdescribed above, or only part of the layer may be scanned andinformation therefor recorded. At a point during scanning, thecumulative dose delivered at each position is compared to a targetcumulative dose for that position. For example, this may be done afterpart of a layer containing that position is scanned, after the entirelayer is scanned, after a set of layers are scanned, or after all layersin an irradiation target are scanned. It is determined (205) if thecurrent cumulative dose matches the target cumulative dose at specificpositions. If the current cumulative dose does match the targetcumulative dose at specific positions, scanning is completed (207) forthose positions. If the current cumulative dose does not match thetarget cumulative dose at specific positions, the scanning system isoperated to compensate for deficiencies in the recorded (e.g., currentcumulative) doses relative to corresponding target cumulative doses forthose positions. For example, if the current cumulative dose does notmatch the target cumulative dose at specific positions, the current inthe scanning magnet may be controlled in order to move (206) the beam soas to deliver additional dose to the specific positions.

As explained above, in some implementations, 100% of the dose may beapplied during a single scan (e.g., a single delivery of particles) of alayer. In that case, more than one scan per layer may not be necessary.In other implementations, less than 100% of the dose may be appliedduring a single scan. In that case, more than one scan per layer will benecessary. To this end, according to the scanning process, for positionsat which dose is applied, if the current cumulative dose at eachposition does not match the target cumulative dose at a correspondingposition, the magnet current is controlled in order to move the beam soas to deliver additional dose to positions that require more dose. Inother words, the layer may be re-scanned any appropriate number of timesuntil the target cumulative dose is reached for all positions of thelayer. In some implementations, in one scan or in multiple scans, theactual delivered dose may exceed 100% of the target cumulative dose.What dose to deliver may be dictated by appropriate medicalprofessionals.

As noted above, the layer may be re-scanned at any appropriate point,e.g., after part of the layer is completed with a current scan, afterthe entire layer is completed with the current scan, after a set oflayers is completed with a scan, or after all layers are completed witha scan. During re-scanning, the process above is repeated until thetarget cumulative dose is reached for all, or some subset of, positionsin the irradiation target. In some implementations, the intensity of theparticle beam may need to be adjusted, e.g., for the last scan. Forexample, if the intensity is set at 25% of the target cumulative dose,but only 20% is delivered at each scan, then a fifth (and possiblysixth) dose will require a lower intensity than 25% in order to reachthe target cumulative dose.

The processes described herein may be used with a single particleaccelerator, and any two or more of the features thereof describedherein may be used with the single particle accelerator. The particleaccelerator may be used in any type of medical or non-medicalapplication. An example of a particle therapy system that may be used isprovided below. Notably, the concepts described herein may be used inother systems not specifically described.

Referring to FIG. 11, an example implementation of a charged particleradiation therapy system 400 includes a beam-producing particleaccelerator 402 (e.g., the particle accelerator of FIGS. 1, 2) having aweight and size small enough to permit it to be mounted on a rotatinggantry 404 with its output directed straight (that is, essentiallydirectly) from the accelerator housing toward a patient 406. Particleaccelerator 402 also includes a scanning system of a type describedherein (e.g., FIGS. 3 to 10).

In some implementations, the steel gantry has two legs 408, 410 mountedfor rotation on two respective bearings 412, 414 that lie on oppositesides of the patient. The accelerator is supported by a steel truss 416that is long enough to span a treatment area 418 in which the patientlies (e.g., twice as long as a tall person, to permit the person to berotated fully within the space with any desired target area of thepatient remaining in the line of the beam) and is attached stably atboth ends to the rotating legs of the gantry.

In some examples, the rotation of the gantry is limited to a range 420of less than 360 degrees, e.g., about 180 degrees, to permit a floor 422to extend from a wall of the vault 424 that houses the therapy systeminto the patient treatment area. The limited rotation range of thegantry also reduces the required thickness of some of the walls (whichare not directly aligned with the beam, e.g., wall 430), which provideradiation shielding of people outside the treatment area. A range of 180degrees of gantry rotation is enough to cover all treatment approachangles, but providing a larger range of travel can be useful. Forexample the range of rotation may be between 180 and 330 degrees andstill provide clearance for the therapy floor space. In otherimplementations, rotation is not limited as described above.

The horizontal rotational axis 432 of the gantry is located nominallyone meter above the floor where the patient and therapist interact withthe therapy system. This floor is positioned about 3 meters above thebottom floor of the therapy system shielded vault. The accelerator canswing under the raised floor for delivery of treatment beams from belowthe rotational axis. The patient couch moves and rotates in asubstantially horizontal plane parallel to the rotational axis of thegantry. The couch can rotate through a range 434 of about 270 degrees inthe horizontal plane with this configuration. This combination of gantryand patient rotational ranges and degrees of freedom allow the therapistto select virtually any approach angle for the beam. If needed, thepatient can be placed on the couch in the opposite orientation and thenall possible angles can be used.

In some implementations, the accelerator uses a synchrocyclotronconfiguration having a high magnetic field superconductingelectromagnetic structure. Because the bend radius of a charged particleof a given kinetic energy is reduced in direct proportion to an increasein the magnetic field applied to it, the high magnetic fieldsuperconducting magnetic structure permits the accelerator to be madesmaller and lighter. The synchrocyclotron uses a magnetic field that isuniform in rotation angle and falls off in strength with increasingradius. Such a field shape can be achieved regardless of the magnitudeof the magnetic field, so in theory there is no upper limit to themagnetic field strength (and therefore the resulting particle energy ata fixed radius) that can be used in a synchrocyclotron.

The synchrocyclotron is supported on the gantry so that the beam isgenerated directly in line with the patient. The gantry permits rotationof the synchrocyclotron about a horizontal rotational axis that containsa point (isocenter 440) within, or near, the patient. The split trussthat is parallel to the rotational axis, supports the synchrocyclotronon both sides.

Because the rotational range of the gantry is limited in some exampleimplementations, a patient support area can be accommodated in a widearea around the isocenter. Because the floor can be extended broadlyaround the isocenter, a patient support table can be positioned to moverelative to and to rotate about a vertical axis 442 through theisocenter so that, by a combination of gantry rotation and table motionand rotation, any angle of beam direction into any part of the patientcan be achieved. In some implementations, the two gantry arms areseparated by more than twice the height of a tall patient, allowing thecouch with patient to rotate and translate in a horizontal plane abovethe raised floor.

Limiting the gantry rotation angle allows for a reduction in thethickness of at least one of the walls surrounding the treatment room.Thick walls, typically constructed of concrete, provide radiationprotection to individuals outside the treatment room. A wall downstreamof a stopping proton beam may be about twice as thick as a wall at theopposite end of the room to provide an equivalent level of protection.Limiting the range of gantry rotation enables the treatment room to besited below earth grade on three sides, while allowing an occupied areaadjacent to the thinnest wall reducing the cost of constructing thetreatment room.

In the example implementation shown in FIG. 11, the superconductingsynchrocyclotron 402 operates with a peak magnetic field in a pole gapof the synchrocyclotron of 8.8 Tesla. The synchrocyclotron produces abeam of protons having an energy of 250 MeV. In some implementations,the synchrocyclotron is a variable-energy machine, and is capable ofoutputting proton beams having different energies. In someimplementations, the synchrocyclotron may produce a beam having a fixedenergy. In some implementations the field strength could be in the rangeof 4 T to 20 T and the proton energy could be in the range of 150 to 300MeV.

The radiation therapy system described in this example is used forproton radiation therapy, but the same principles and details can beapplied in analogous systems for use in heavy ion (ion) treatmentsystems.

As shown in FIGS. 1, 2, 12, 13, and 14, an example synchrocyclotron 10(e.g., 402 in FIG. 11) includes a magnet system 122 that contains aparticle source 190, a radiofrequency drive system 191, and a beamextraction system. In this example, the magnetic field established bythe magnet system has a shape appropriate to maintain focus of acontained proton beam using a combination of a split pair of annularsuperconducting coils 140, 142 and a pair of shaped ferromagnetic (e.g.,low carbon steel) pole faces 144, 146.

The two superconducting magnet coils are centered on a common axis andare spaced apart along the axis. The coils may be formed by ofNb₃Sn-based superconducting 0.8 mm diameter strands (that initiallycomprise a niobium-tin core surrounded by a copper sheath) deployed in atwisted cable-in-channel conductor geometry. After seven individualstrands are cabled together, they are heated to cause a reaction thatforms the final (brittle) superconducting material of the wire. Afterthe material has been reacted, the wires are soldered into the copperchannel (outer dimensions 3.18×2.54 mm and inner dimensions 2.08×2.08mm) and covered with insulation (in this example, a woven fiberglassmaterial). The copper channel containing the wires is then wound in acoil having a rectangular cross-section. The wound coil is then vacuumimpregnated with an epoxy compound. The finished coils are mounted on anannular stainless steel reverse bobbin. Heater blankets may be placed atintervals in the layers of the windings to protect the assembly in theevent of a magnet quench.

The entire coil can then be covered with copper sheets to providethermal conductivity and mechanical stability and then contained in anadditional layer of epoxy. The precompression of the coil can beprovided by heating the stainless steel reverse bobbin and fitting thecoils within the reverse bobbin. The reverse bobbin inner diameter ischosen so that when the entire mass is cooled to 4 K, the reverse bobbinstays in contact with the coil and provides some compression. Heatingthe stainless steel reverse bobbin to approximately 50 degrees C. andfitting coils at a temperature of 100 degrees Kelvin can achieve this.

The geometry of the coil is maintained by mounting the coils in a“reverse” rectangular bobbin to exert a restorative force that worksagainst the distorting force produced when the coils are energized. Asshown in FIG. 13, in some implementations, coil position is maintainedrelative to corresponding magnet pole pieces and the cryostat using aset of warm-to-cold support straps 402, 404, 406. Supporting the coldmass with thin straps reduces the heat leakage imparted to the cold massby the rigid support system. The straps are arranged to withstand thevarying gravitational force on the coil as the magnet rotates on boardthe gantry. They withstand the combined effects of gravity and the largede-centering force realized by the coil when it is perturbed from aperfectly symmetric position relative to the magnet yoke. Additionally,the links act to reduce dynamic forces imparted on the coil as thegantry accelerates and decelerates when its position is changed. Eachwarm-to-cold support may include one S2 fiberglass link and one carbonfiber link. The carbon fiber link is supported across pins between thewarm yoke and an intermediate temperature (50-70 K), and the S2fiberglass link 408 is supported across the intermediate temperature pinand a pin attached to the cold mass. Each pin may be made of highstrength stainless steel.

Referring to FIG. 1, the field strength profile as a function of radiusis determined largely by choice of coil geometry and pole face shape.The pole faces 144, 146 of the permeable yoke material can be contouredto fine tune the shape of the magnetic field to ensure that the particlebeam remains focused during acceleration.

The superconducting coils are maintained at temperatures near absolutezero (e.g., about 4 degrees Kelvin) by enclosing the coil assembly (thecoils and the bobbin) inside an evacuated annular aluminum or stainlesssteel cryostatic chamber 170 (the cryostat) that provides a free spacearound the coil structure, except at a limited set of support points171, 173. In an alternate version (e.g., FIG. 2) the outer wall of thecryostat may be made of low carbon steel to provide an additional returnflux path for the magnetic field.

In some implementations, the temperature near absolute zero is achievedand maintained using one single-stage Gifford-McMahon cryo-cooler andthree two-stage Gifford McMahon cryo-coolers. Each two stage cryo-coolerhas a second stage cold end attached to a condenser that recondensesHelium vapor into liquid Helium. In some implementations, thetemperature near absolute zero is achieved and maintained using acooling channel (not shown) containing liquid helium, which is formedinside a superconducting coil support structure (e.g., the reversebobbin), and which contains a thermal connection between the liquidhelium in the channel and the corresponding superconducting coil. Anexample of a liquid helium cooling system of the type described above,and that may be used is described in U.S. patent application Ser. No.13/148,000 (Begg et al.).

In some implementations, the coil assembly and cryostatic chambers aremounted within and fully enclosed by two halves 181, 183 of apillbox-shaped magnet yoke 100. The yoke 100 provides a path for thereturn magnetic field flux 184 and magnetically shields the volume 186between the pole faces 144, 146 to prevent external magnetic influencesfrom perturbing the shape of the magnetic field within that volume. Theyoke also serves to decrease the stray magnetic field in the vicinity ofthe accelerator. In other implementations, the coil assembly andcryostatic chambers are mounted within and fully enclosed by anon-magnetic enclosure, and the path for return magnetic field flux isimplemented using an active return system, an example of which isdescribed above.

As shown in FIGS. 1 and 15, the synchrocyclotron includes a particlesource 190 of a Penning ion gauge geometry located near the geometriccenter 192 of the magnet structure. The particle source may be asdescribed below, or the particle source may be of the type described inU.S. patent application Ser. No. 11/948,662 incorporated herein byreference.

Particle source 190 is fed from a supply 399 of hydrogen through a gasline 393 and tube 394 that delivers gaseous hydrogen. Electric cables294 carry an electric current from a current source to stimulateelectron discharge from cathodes 392, 390 that are aligned with themagnetic field.

In this example, the discharged electrons ionize the gas exiting througha small hole from tube 394 to create a supply of positive ions (protons)for acceleration by one semicircular (dee-shaped) radio-frequency platethat spans half of the space enclosed by the magnet structure and onedummy dee plate. In the case of an interrupted particle source (anexample of which is described in U.S. patent application Ser. No.11/948,662), all (or a substantial part, e.g., a majority) of the tubecontaining plasma is removed at the acceleration region.

As shown in FIG. 16, the dee plate 500 is a hollow metal structure thathas two semicircular surfaces 503, 505 that enclose a space 507 in whichthe protons are accelerated during half of their rotation around thespace enclosed by the magnet structure. A duct 509 opening into thespace 507 extends through the enclosure (e.g., the yoke or polepiece(s)) to an external location from which a vacuum pump can beattached to evacuate the space 507 and the rest of the space within avacuum chamber in which the acceleration takes place. The dummy dee 502comprises a rectangular metal ring that is spaced near to the exposedrim of the dee plate. The dummy dee is grounded to the vacuum chamberand magnet yoke. The dee plate 500 is driven by a radio-frequency signalthat is applied at the end of a radio-frequency transmission line toimpart an electric field in the space 507. The radio frequency electricfield is made to vary in time as the accelerated particle beam increasesin distance from the geometric center. The radio frequency electricfield may be controlled in the manner described in U.S. patentapplication Ser. No. 11/948,359, entitled “Matching A Resonant FrequencyOf A Resonant Cavity To A Frequency Of An Input Voltage”, the contentsof which are incorporated herein by reference.

For the beam emerging from the centrally located particle source toclear the particle source structure as it begins to spiral outward, alarge voltage difference may be applied across the radio frequencyplates. 20,000 Volts is applied across the radio frequency plates. Insome versions from 8,000 to 20,000 Volts may be applied across the radiofrequency plates. To reduce the power required to drive this largevoltage, the magnet structure is arranged to reduce the capacitancebetween the radio frequency plates and ground. This may be done byforming holes with sufficient clearance from the radio frequencystructures through the outer yoke and the cryostat housing and makingsufficient space between the magnet pole faces.

The high voltage alternating potential that drives the dee plate has afrequency that is swept downward during the accelerating cycle toaccount for the increasing relativistic mass of the protons and thedecreasing magnetic field. The dummy dee does not require a hollowsemi-cylindrical structure as it is at ground potential along with thevacuum chamber walls. Other plate arrangements could be used such asmore than one pair of accelerating electrodes driven with differentelectrical phases or multiples of the fundamental frequency. The RFstructure can be tuned to keep the Q high during the required frequencysweep by using, for example, a rotating capacitor having intermeshingrotating and stationary blades. During each meshing of the blades, thecapacitance increases, thus lowering the resonant frequency of the RFstructure. The blades can be shaped to create a precise frequency sweeprequired. A drive motor for the rotating condenser can be phase lockedto the RF generator for precise control. One bunch of particles may beaccelerated during each meshing of the blades of the rotating condenser.

The vacuum chamber in which the acceleration occurs is a generallycylindrical container that is thinner in the center and thicker at therim. The vacuum chamber encloses the RF plates and the particle sourceand is evacuated by a vacuum pump. Maintaining a high vacuum reduces thechances that accelerating ions are not lost to collisions with gasmolecules and enables the RF voltage to be kept at a higher levelwithout arcing to ground.

Protons (or other ions) traverse a generally spiral orbital pathbeginning at the particle source. In half of each loop of the spiralpath, the protons gain energy as they pass through the RF electricfield. As the protons gain energy, the radius of the central orbit ofeach successive loop of their spiral path is larger than the prior loopuntil the loop radius reaches the maximum radius of the pole face. Atthat location a magnetic and electric field perturbation directs protonsinto an area where the magnetic field rapidly decreases, and the protonsdepart the area of the high magnetic field and are directed through anevacuated tube, referred to herein as the extraction channel, to exitthe synchrocyclotron. A magnetic regenerator may be used to change themagnetic field perturbation to direct the protons. The protons exitingwill tend to disperse as they enter an area of markedly decreasedmagnetic field that exists in the room around the synchrocyclotron. Beamshaping elements 507, 509 in the extraction channel 138 (FIG. 13)redirect the protons so that they stay in a straight beam of limitedspatial extent.

As the beam exits the extraction channel it is passed through a beamformation system 525 (FIG. 13), which may include a scanning system ofthe type described herein. Beam formation system 525 may be used inconjunction with an inner gantry that controls application of the beam.

Stray magnetic fields exiting from the synchrocyclotron may be limitedby both a magnet yoke (which also serves as a shield) and a separatemagnetic shield 514 (e.g., FIG. 1). The separate magnetic shieldincludes of a layer 517 of ferromagnetic material (e.g., steel or iron)that encloses the pillbox yoke, separated by a space 516. Thisconfiguration that includes a sandwich of a yoke, a space, and a shieldachieves adequate shielding for a given leakage magnetic field at lowerweight. As described above, in some implementations, an active returnsystem may be used in place of, or to augment, the operation of themagnetic yoke and shield.

Referring to FIG. 11, the gantry allows the synchrocyclotron to berotated about a horizontal rotational axis 432. The truss structure 416has two generally parallel spans 480, 482. The synchrocyclotron iscradled between the spans about midway between the legs. The gantry isbalanced for rotation about the bearings using counterweights 622, 624mounted on ends of the legs opposite the truss.

The gantry is driven to rotate by an electric motor mounted to one orboth of the gantry legs and connected to the bearing housings by drivegears. The rotational position of the gantry is derived from signalsprovided by shaft angle encoders incorporated into the gantry drivemotors and the drive gears.

At the location at which the ion beam exits the synchrocyclotron, thebeam formation system 525 acts on the ion beam to give it propertiessuitable for patient treatment. For example, the beam may be spread andits depth of penetration varied to provide uniform radiation across agiven target volume. The beam formation system may include activescanning elements as described herein.

All of the active systems of the synchrocyclotron (the current drivensuperconducting coils, the RF-driven plates, the vacuum pumps for thevacuum acceleration chamber and for the superconducting coil coolingchamber, the current driven particle source, the hydrogen gas source,and the RF plate coolers, for example), may be controlled by appropriatesynchrocyclotron control electronics (not shown), which may include,e.g., one or more processing devices executing instructions fromnon-transitory memory to effect control.

As explained above, referring to system 602 of FIG. 17, a beam-producingparticle accelerator, in this case synchrocyclotron 604 (which mayinclude any and all features described herein), may be mounted onrotating gantry 605. Rotating gantry 605 is of the type describedherein, and can angularly rotate around patient support 606. Thisfeature enables synchrocyclotron 604 to provide a particle beamessentially directly to the patient from various angles. For example, asin FIG. 17, if synchrocyclotron 604 is above patient support 606, theparticle beam may be directed downwards toward the patient.Alternatively, if synchrocyclotron 604 is below patient support 606, theparticle beam may be directed upwards toward the patient. The particlebeam is applied essentially directly to the patient in the sense that anintermediary beam routing mechanism is not required. A routingmechanism, in this context, is different from a shaping or sizingmechanism in that a shaping or sizing mechanism does not re-route thebeam, but rather sizes and/or shapes the beam while maintaining the samegeneral trajectory of the beam.

Further details regarding an example implementation of the foregoingsystem may be found in U.S. Pat. No. 7,728,311, filed on Nov. 16, 2006and entitled “Charged Particle Radiation Therapy”, and in U.S. patentapplication Ser. No. 12/275,103, filed on Nov. 20, 2008 and entitled“Inner Gantry”. The contents of U.S. Pat. No. 7,728,311 and in U.S.patent application Ser. No. 12/275,103 are hereby incorporated byreference into this disclosure. In some implementations, thesynchrocyclotron may be a variable-energy device, such as that describedin U.S. patent application Ser. No. 13/916,401, filed on Jun. 12, 2013,the contents of which are incorporated herein by reference.

VARIABLE-ENERGY PARTICLE ACCELERATOR

The particle accelerator used in the example particle therapy systemsand example scanning systems described herein may be a variable-energyparticle accelerator, an example of which is described below

The energy of an extracted particle beam (the particle beam output fromthe accelerator) can affect the use of the particle beam duringtreatment. In some machines, the energy of the particle beam (orparticles in the particle beam) does not increase after extraction.However, the energy may be reduced based on treatment needs after theextraction and before the treatment. Referring to FIG. 18, an exampletreatment system 910 includes an accelerator 912, e.g., asynchrocyclotron, from which a particle (e.g., proton) beam 914 having avariable energy is extracted to irradiate a target volume 924 of a body922. Optionally, one or more additional devices, such as a scanning unit916 or a scattering unit 916, one or more monitoring units 918, and anenergy degrader 920, are placed along the irradiation direction 928. Thedevices intercept the cross-section of the extracted beam 914 and alterone or more properties of the extracted beam for the treatment.

A target volume to be irradiated (an irradiation target) by a particlebeam for treatment typically has a three-dimensional configuration. Insome examples, to carry-out the treatment, the target volume is dividedinto layers along the irradiation direction of the particle beam so thatthe irradiation can be done on a layer-by-layer basis. For certain typesof particles, such as protons, the penetration depth (or which layer thebeam reaches) within the target volume is largely determined by theenergy of the particle beam. A particle beam of a given energy does notreach substantially beyond a corresponding penetration depth for thatenergy. To move the beam irradiation from one layer to another layer ofthe target volume, the energy of the particle beam is changed.

In the example shown in FIG. 18, the target volume 924 is divided intonine layers 926 a-926 i along the irradiation direction 928. In anexample process, the irradiation starts from the deepest layer 926 i,one layer at a time, gradually to the shallower layers and finishes withthe shallowest layer 926 a. Before application to the body 922, theenergy of the particle beam 914 is controlled to be at a level to allowthe particle beam to stop at a desired layer, e.g., the layer 926 d,without substantially penetrating further into the body or the targetvolume, e.g., the layers 926 e-926 i or deeper into the body. In someexamples, the desired energy of the particle beam 914 decreases as thetreatment layer becomes shallower relative to the particle acceleration.In some examples, the beam energy difference for treating adjacentlayers of the target volume 924 is about 3 MeV to about 100 MeV, e.g.,about 10 MeV to about 80 MeV, although other differences may also bepossible, depending on, e.g., the thickness of the layers and theproperties of the beam.

The energy variation for treating different layers of the target volume924 can be performed at the accelerator 912 (e.g., the accelerator canvary the energy) so that, in some implementations, no additional energyvariation is required after the particle beam is extracted from theaccelerator 912. So, the optional energy degrader 920 in the treatmentsystem 10 may be eliminated from the system. In some implementations,the accelerator 912 can output particle beams having an energy thatvaries between about 100 MeV and about 300 MeV, e.g., between about 115MeV and about 250 MeV. The variation can be continuous ornon-continuous, e.g., one step at a time. In some implementations, thevariation, continuous or non-continuous, can take place at a relativelyhigh rate, e.g., up to about 50 MeV per second or up to about 20 MeV persecond. Non-continuous variation can take place one step at a time witha step size of about 10 MeV to about 90 MeV.

When irradiation is complete in one layer, the accelerator 912 can varythe energy of the particle beam for irradiating a next layer, e.g.,within several seconds or within less than one second. In someimplementations, the treatment of the target volume 924 can be continuedwithout substantial interruption or even without any interruption. Insome situations, the step size of the non-continuous energy variation isselected to correspond to the energy difference needed for irradiatingtwo adjacent layers of the target volume 924. For example, the step sizecan be the same as, or a fraction of, the energy difference.

In some implementations, the accelerator 912 and the degrader 920collectively vary the energy of the beam 914. For example, theaccelerator 912 provides a coarse adjustment and the degrader 920provides a fine adjustment or vice versa. In this example, theaccelerator 912 can output the particle beam that varies energy with avariation step of about 10-80 MeV, and the degrader 920 adjusts (e.g.,reduces) the energy of the beam at a variation step of about 2-10 MeV.

The reduced use (or absence) of the energy degrader, such as a rangemodulator, may help to maintain properties and quality of the outputbeam from the accelerator, e.g., beam intensity. The control of theparticle beam can be performed at the accelerator. Side effects, e.g.,from neutrons generated when the particle beam passes the degrader 920can be reduced or eliminated.

The energy of the particle beam 914 may be adjusted to treat anothertarget volume 930 in another body or body part 922′ after completingtreatment in target volume 924. The target volumes 924, 930 may be inthe same body (or patient), or in different patients. It is possiblethat the depth D of the target volume 930 from a surface of body 922′ isdifferent from that of the target volume 924. Although some energyadjustment may be performed by the degrader 920, the degrader 912 mayonly reduce the beam energy and not increase the beam energy.

In this regard, in some cases, the beam energy required for treatingtarget volume 930 is greater than the beam energy required to treattarget volume 924. In such cases, the accelerator 912 may increase theoutput beam energy after treating the target volume 924 and beforetreating the target volume 930. In other cases, the beam energy requiredfor treating target volume 930 is less than the beam energy required totreat target volume 924. Although the degrader 920 can reduce theenergy, the accelerator 912 can be adjusted to output a lower beamenergy to reduce or eliminate the use of the degrader 920. The divisionof the target volumes 924, 930 into layers can be different or the same.The target volume 930 can be treated similarly on a layer by layer basisto the treatment of the target volume 924.

The treatment of the different target volumes 924, 930 on the samepatient may be substantially continuous, e.g., with the stop timebetween the two volumes being no longer than about 30 minutes or less,e.g., 25 minutes or less, 20 minutes or less, 15 minutes or less, 10minutes or less, 5 minutes or less, or 1 minute or less. As explainedherein, the accelerator 912 can be mounted on a movable gantry and themovement of the gantry can move the accelerator to aim at differenttarget volumes. In some situations, the accelerator 912 can complete theenergy adjustment of the output beam 914 during the time the treatmentsystem makes adjustment (such as moving the gantry) after completing thetreatment of the target volume 924 and before starting treating thetarget volume 930. After the alignment of the accelerator and the targetvolume 930, the treatment can begin with the adjusted, desired beamenergy. Beam energy adjustment for different patients can also becompleted relatively efficiently. In some examples, all adjustments,including increasing/reducing beam energy and/or moving the gantry aredone within about 30 minutes, e.g., within about 25 minutes, withinabout 20 minutes, within about 15 minutes, within about 10 minutes orwithin about 5 minutes.

In the same layer of a target volume, an irradiation dose may be appliedby moving the beam across the two-dimensional surface of the layer(which is sometimes called scanning beam) using a scanning unit 916.Alternatively, the layer can be irradiated by passing the extracted beamthrough one or more scatterers of the scattering unit 16 (which issometimes called scattering beam).

Beam properties, such as energy and intensity, can be selected before atreatment or can be adjusted during the treatment by controlling theaccelerator 912 and/or other devices, such as the scanningunit/scatterer(s) 916, the degrader 920, and others not shown in thefigures. In example implementations, system 910 includes a controller932, such as a computer, in communication with one or more devices inthe system. Control can be based on results of the monitoring performedby the one or more monitors 918, e.g., monitoring of the beam intensity,dose, beam location in the target volume, etc. Although the monitors 918are shown to be between the device 916 and the degrader 920, one or moremonitors can be placed at other appropriate locations along the beamirradiation path. Controller 932 can also store a treatment plan for oneor more target volumes (for the same patient and/or different patients).The treatment plan can be determined before the treatment starts and caninclude parameters, such as the shape of the target volume, the numberof irradiation layers, the irradiation dose for each layer, the numberof times each layer is irradiated, etc. The adjustment of a beamproperty within the system 910 can be performed based on the treatmentplan. Additional adjustment can be made during the treatment, e.g., whendeviation from the treatment plan is detected.

In some implementations, the accelerator 912 is configured to vary theenergy of the output particle beam by varying the magnetic field inwhich the particle beam is accelerated. In an example implementation,one or more sets of coils receives variable electrical current toproduce a variable magnetic field in the cavity. In some examples, oneset of coils receives a fixed electrical current, while one or moreother sets of coils receives a variable current so that the totalcurrent received by the coil sets varies. In some implementations, allsets of coils are superconducting. In other implementations, some setsof coils, such as the set for the fixed electrical current, aresuperconducting, while other sets of coils, such as the one or more setsfor the variable current, are non-superconducting. In some examples, allsets of coils are non-superconducting.

Generally, the magnitude of the magnetic field is scalable with themagnitude of the electrical current. Adjusting the total electriccurrent of the coils in a predetermined range can generate a magneticfield that varies in a corresponding, predetermined range. In someexamples, a continuous adjustment of the electrical current can lead toa continuous variation of the magnetic field and a continuous variationof the output beam energy. Alternatively, when the electrical currentapplied to the coils is adjusted in a non-continuous, step-wise manner,the magnetic field and the output beam energy also varies accordingly ina non-continuous (step-wise) manner. The scaling of the magnetic fieldto the current can allow the variation of the beam energy to be carriedout relatively precisely, although sometimes minor adjustment other thanthe input current may be performed.

In some implementations, to output particle beams having a variableenergy, the accelerator 912 is configured to apply RF voltages thatsweep over different ranges of frequencies, with each rangecorresponding to a different output beam energy. For example, if theaccelerator 912 is configured to produce three different output beamenergies, the RF voltage is capable of sweeping over three differentranges of frequencies. In another example, corresponding to continuousbeam energy variations, the RF voltage sweeps over frequency ranges thatcontinuously change. The different frequency ranges may have differentlower frequency and/or upper frequency boundaries.

The extraction channel may be configured to accommodate the range ofdifferent energies produced by the variable-energy particle accelerator.For example the extraction channel may be large enough to support thehighest and lowest energies produced by the particle accelerator. Thatis, the extraction channel may be sized or otherwise configured toreceive and to transmit particles within that range of energies.Particle beams having different energies can be extracted from theaccelerator 912 without altering the features of the regenerator that isused for extracting particle beams having a single energy. In otherimplementations, to accommodate the variable particle energy, theregenerator can be moved to disturb (e.g., change) different particleorbits in the manner described above and/or iron rods (magnetic shims)can be added or removed to change the magnetic field bump provided bythe regenerator. More specifically, different particle energies willtypically be at different particle orbits within the cavity. By movingthe regenerator, it is possible to intercept a particle orbit at aspecified energy and thereby provide the correct perturbation of thatorbit so that particles at the specified energy reach the extractionchannel. In some implementations, movement of the regenerator (and/oraddition/removal of magnetic shims) is performed in real-time to matchreal-time changes in the particle beam energy output by the accelerator.In other implementations, particle energy is adjusted on a per-treatmentbasis, and movement of the regenerator (and/or addition/removal ofmagnetic shims) is performed in advance of the treatment. In eithercase, movement of the regenerator (and/or addition/removal of magneticshims) may be computer controlled. For example, a computer may controlone or more motors that effect movement of the regenerator and/ormagnetic shims.

In some implementations, the regenerator is implemented using one ormore magnetic shims that are controllable to move to the appropriatelocation(s).

As an example, table 1 shows three example energy levels at whichexample accelerator 912 can output particle beams. The correspondingparameters for producing the three energy levels are also listed. Inthis regard, the magnet current refers to the total electrical currentapplied to the one or more coil sets in the accelerator 912; the maximumand minimum frequencies define the ranges in which the RF voltagesweeps; and “r” is the radial distance of a location to a center of thecavity in which the particles are accelerated.

TABLE 1 Examples of beam energies and respective parameters. MagneticMagnetic Beam Magnet Maximum Minimum Field at Field at Energy CurrentFrequency Frequency r = 0 mm r = 298 mm (MeV) (Amps) (MHz) (MHz) (Tesla)(Tesla) 250 1990 132 99 8.7 8.2 235 1920 128 97 8.4 8.0 211 1760 120 937.9 7.5

Details that may be included in an example particle accelerator thatproduces charged particles having variable energies are described below.The accelerator can be a synchrocyclotron and the particles may beprotons. The particles may be output as pulsed beams. The energy of thebeam output from the particle accelerator can be varied during thetreatment of one target volume in a patient, or between treatments ofdifferent target volumes of the same patient or different patients. Insome implementations, settings of the accelerator are changed to varythe beam energy when no beam (or particles) is output from theaccelerator. The energy variation can be continuous or non-continuousover a desired range.

Referring to the example shown in FIG. 1, the particle accelerator,which may be a variable-energy particle accelerator like accelerator 912described above, may be configured to output particle beams that have avariable energy. The range of the variable energy can have an upperboundary that is about 200 MeV to about 300 MeV or higher, e.g., 200MeV, about 205 MeV, about 210 MeV, about 215 MeV, about 220 MeV, about225 MeV, about 230 MeV, about 235 MeV, about 240 MeV, about 245 MeV,about 250 MeV, about 255 MeV, about 260 MeV, about 265 MeV, about 270MeV, about 275 MeV, about 280 MeV, about 285 MeV, about 290 MeV, about295 MeV, or about 300 MeV or higher. The range can also have a lowerboundary that is about 100 MeV or lower to about 200 MeV, e.g., about100 MeV or lower, about 105 MeV, about 110 MeV, about 115 MeV, about 120MeV, about 125 MeV, about 130 MeV, about 135 MeV, about 140 MeV, about145 MeV, about 150 MeV, about 155 MeV, about 160 MeV, about 165 MeV,about 170 MeV, about 175 MeV, about 180 MeV, about 185 MeV, about 190MeV, about 195 MeV, about 200 MeV.

In some examples, the variation is non-continuous and the variation stepcan have a size of about 10 MeV or lower, about 15 MeV, about 20 MeV,about 25 MeV, about 30 MeV, about 35 MeV, about 40 MeV, about 45 MeV,about 50 MeV, about 55 MeV, about 60 MeV, about 65 MeV, about 70 MeV,about 75 MeV, or about 80 MeV or higher. Varying the energy by one stepsize can take no more than 30 minutes, e.g., about 25 minutes or less,about 20 minutes or less, about 15 minutes or less, about 10 minutes orless, about 5 minutes or less, about 1 minute or less, or about 30seconds or less. In other examples, the variation is continuous and theaccelerator can adjust the energy of the particle beam at a relativelyhigh rate, e.g., up to about 50 MeV per second, up to about 45 MeV persecond, up to about 40 MeV per second, up to about 35 MeV per second, upto about 30 MeV per second, up to about 25 MeV per second, up to about20 MeV per second, up to about 15 MeV per second, or up to about 10 MeVper second. The accelerator can be configured to adjust the particleenergy both continuously and non-continuously. For example, acombination of the continuous and non-continuous variation can be usedin a treatment of one target volume or in treatments of different targetvolumes. Flexible treatment planning and flexible treatment can beachieved.

A particle accelerator that outputs a particle beam having a variableenergy can provide accuracy in irradiation treatment and reduce thenumber of additional devices (other than the accelerator) used for thetreatment. For example, the use of degraders for changing the energy ofan output particle beam may be reduced or eliminated for all or part ofthe treatment. The properties of the particle beam, such as intensity,focus, etc. can be controlled at the particle accelerator and theparticle beam can reach the target volume without substantialdisturbance from the additional devices. The relatively high variationrate of the beam energy can reduce treatment time and allow forefficient use of the treatment system.

In some implementations, the accelerator, such as the synchrocyclotronof FIG. 1, accelerates particles or particle beams to variable energylevels by varying the magnetic field in the accelerator, which can beachieved by varying the electrical current applied to coils forgenerating the magnetic field. As explained above, an examplesynchrocyclotron (e.g., the synchrocyclotron of FIG. 1) includes amagnet system that contains a particle source, a radiofrequency drivesystem, and a beam extraction system. FIG. 19 shows an example of amagnet system that may be used in a variable-energy accelerator. In thisexample implementation, the magnetic field established by the magnetsystem 1012 can vary by about 5% to about 35% of a maximum value of themagnetic field that two sets of coils 40 a and 40 b, and 42 a and 42 bare capable of generating. The magnetic field established by the magnetsystem has a shape appropriate to maintain focus of a contained protonbeam using a combination of the two sets of coils and a pair of shapedferromagnetic (e.g., low carbon steel) structures, examples of which areprovided above.

Each set of coils may be a split pair of annular coils to receiveelectrical current. In some situations, both sets of coils aresuperconducting. In other situations, only one set of the coils issuperconducting and the other set is non-superconducting or normalconducting (also discussed further below). It is also possible that bothsets of coils are non-superconducting. Suitable superconductingmaterials for use in the coils include niobium-3 tin (Nb3Sn) and/orniobium-titanium. Other normal conducting materials can include copper.Examples of the coil set constructions are described further below.

The two sets of coils can be electrically connected serially or inparallel. In some implementations, the total electrical current receivedby the two sets of coils can include about 2 million ampere turns toabout 10 million ampere turns, e.g., about 2.5 to about 7.5 millionampere turns or about 3.75 million ampere turns to about 5 millionampere turns. In some examples, one set of coils is configured toreceive a fixed (or constant) portion of the total variable electricalcurrent, while the other set of coils is configured to receive avariable portion of the total electrical current. The total electricalcurrent of the two coil sets varies with the variation of the current inone coil set. In other situations, the electrical current applied toboth sets of coils can vary. The variable total current in the two setsof coils can generate a magnetic field having a variable magnitude,which in turn varies the acceleration pathways of the particles andproduces particles having variable energies.

Generally, the magnitude of the magnetic field generated by the coil(s)is scalable to the magnitude of the total electrical current applied tothe coil(s). Based on the scalability, in some implementations, linearvariation of the magnetic field strength can be achieved by linearlychanging the total current of the coil sets. The total current can beadjusted at a relatively high rate that leads to a relatively high-rateadjustment of the magnetic field and the beam energy.

In the example reflected in Table 1 above, the ratio between values ofthe current and the magnetic field at the geometric center of the coilrings is: 1990:8.7 (approximately 228.7:1); 1920:8.4 (approximately228.6:1); 1760:7.9 (approximately 222.8:1). Accordingly, adjusting themagnitude of the total current applied to a superconducting coil(s) canproportionally (based on the ratio) adjust the magnitude of the magneticfield.

The scalability of the magnetic field to the total electrical current inthe example of Table 1 is also shown in the plot of FIG. 20, where BZ isthe magnetic field along the Z direction; and R is the radial distancemeasured from a geometric center of the coil rings along a directionperpendicular to the Z direction. The magnetic field has the highestvalue at the geometric center, and decreases as the distance Rincreases. The curves 1035, 1037 represent the magnetic field generatedby the same coil sets receiving different total electrical current: 1760Amperes and 1990 Amperes, respectively. The corresponding energies ofthe extracted particles are 211 MeV and 250 MeV, respectively. The twocurves 1035, 1037 have substantially the same shape and the differentparts of the curves 1035, 1037 are substantially parallel. As a result,either the curve 1035 or the curve 1037 can be linearly shifted tosubstantially match the other curve, indicating that the magnetic fieldis scalable to the total electrical current applied to the coil sets.

In some implementations, the scalability of the magnetic field to thetotal electrical current may not be perfect. For example, the ratiobetween the magnetic field and the current calculated based on theexample shown in table 1 is not constant. Also, as shown in FIG. 21, thelinear shift of one curve may not perfectly match the other curve. Insome implementations, the total current is applied to the coil setsunder the assumption of perfect scalability. The target magnetic field(under the assumption of perfect scalability) can be generated byadditionally altering the features, e.g., geometry, of the coils tocounteract the imperfection in the scalability. As one example,ferromagnetic (e.g., iron) rods (magnetic shims) can be inserted orremoved from one or both of the magnetic structures (e.g., yokes, polepieces, and the like). The features of the coils can be altered at arelatively high rate so that the rate of the magnetic field adjustmentis not substantially affected as compared to the situation in which thescalability is perfect and only the electrical current needs to beadjusted. In the example of iron rods, the rods can be added or removedat the time scale of seconds or minutes, e.g., within 5 minutes, within1 minute, less than 30 seconds, or less than 1 second.

In some implementations, settings of the accelerator, such as thecurrent applied to the coil sets, can be chosen based on the substantialscalability of the magnetic field to the total electrical current in thecoil sets.

Generally, to produce the total current that varies within a desiredrange, any appropriate combination of current applied to the two coilsets can be used. In an example, the coil set 42 a, 42 b can beconfigured to receive a fixed electrical current corresponding to alower boundary of a desired range of the magnetic field. In the exampleshown in table 1, the fixed electrical current is 1760 Amperes. Inaddition, the coil set 40 a, 40 b can be configured to receive avariable electrical current having an upper boundary corresponding to adifference between an upper boundary and a lower boundary of the desiredrange of the magnetic field. In the example shown in table 1, the coilset 40 a, 40 b is configured to receive electrical current that variesbetween 0 Ampere and 230 Amperes.

In another example, the coil set 42 a, 42 b can be configured to receivea fixed electrical current corresponding to an upper boundary of adesired range of the magnetic field. In the example shown in table 1,the fixed current is 1990 Amperes. In addition, the coil set 40 a, 40 bcan be configured to receive a variable electrical current having anupper boundary corresponding to a difference between a lower boundaryand an upper boundary of the desired range of the magnetic field. In theexample shown in table 1, the coil set 40 a, 40 b is configured toreceive electrical current that varies between −230 Ampere and 0 Ampere.

The total variable magnetic field generated by the variable totalcurrent for accelerating the particles can have a maximum magnitudegreater than 4 Tesla, e.g., greater than 5 Tesla, greater than 6 Tesla,greater than 7 Tesla, greater than 8 Tesla, greater than 9 Tesla, orgreater than 10 Tesla, and up to about 20 Tesla or higher, e.g., up toabout 18 Tesla, up to about 15 Tesla, or up to about 12 Tesla. In someimplementations, variation of the total current in the coil sets canvary the magnetic field by about 0.2 Tesla to about 4.2 Tesla or more,e.g., about 0.2 Tesla to about 1.4 Tesla or about 0.6 Tesla to about 4.2Tesla. In some situations, the amount of variation of the magnetic fieldcan be proportional to the maximum magnitude.

FIG. 21 shows an example RF structure for sweeping the voltage on thedee plate 500 over an RF frequency range for each energy level of theparticle beam, and for varying the frequency range when the particlebeam energy is varied. The semicircular surfaces 503, 505 of the deeplate 500 are connected to an inner conductor 1300 and housed in anouter conductor 1302. The high voltage is applied to the dee plate 500from a power source (not shown, e.g., an oscillating voltage input)through a power coupling device 1304 that couples the power source tothe inner conductor. In some implementations, the coupling device 1304is positioned on the inner conductor 1300 to provide power transfer fromthe power source to the dee plate 500. In addition, the dee plate 500 iscoupled to variable reactive elements 1306, 1308 to perform the RFfrequency sweep for each particle energy level, and to change the RFfrequency range for different particle energy levels.

The variable reactive element 1306 can be a rotating capacitor that hasmultiple blades 1310 rotatable by a motor (not shown). By meshing orunmeshing the blades 1310 during each cycle of RF sweeping, thecapacitance of the RF structure changes, which in turn changes theresonant frequency of the RF structure. In some implementations, duringeach quarter cycle of the motor, the blades 1310 mesh with the eachother. The capacitance of the RF structure increases and the resonantfrequency decreases. The process reverses as the blades 1310 unmesh. Asa result, the power required to generate the high voltage applied to thedee plate 103 and necessary to accelerate the beam can be reduced by alarge factor. In some implementations, the shape of the blades 1310 ismachined to form the required dependence of resonant frequency on time.

The RF frequency generation is synchronized with the blade rotation bysensing the phase of the RF voltage in the resonator, keeping thealternating voltage on the dee plates close to the resonant frequency ofthe RF cavity. (The dummy dee is grounded and is not shown in FIG. 21).

The variable reactive element 1308 can be a capacitor formed by a plate1312 and a surface 1316 of the inner conductor 1300. The plate 1312 ismovable along a direction 1314 towards or away from the surface 1316.The capacitance of the capacitor changes as the distance D between theplate 1312 and the surface 1316 changes. For each frequency range to beswept for one particle energy, the distance D is at a set value, and tochange the frequency range, the plate 1312 is moved corresponding to thechange in the energy of the output beam.

In some implementations, the inner and outer conductors 1300, 1302 areformed of a metallic material, such as copper, aluminum, or silver. Theblades 1310 and the plate 1312 can also be formed of the same ordifferent metallic materials as the conductors 1300, 1302. The couplingdevice 1304 can be an electrical conductor. The variable reactiveelements 1306, 1308 can have other forms and can couple to the dee plate100 in other ways to perform the RF frequency sweep and the frequencyrange alteration. In some implementations, a single variable reactiveelement can be configured to perform the functions of both the variablereactive elements 1306, 1308. In other implementations, more than twovariable reactive elements can be used.

The control of the gantry, the patient support, the active beam shapingelements, and the synchrocyclotron to perform a therapy session isachieved by appropriate therapy control electronics (not shown).

Control of the particle therapy system described herein and its variousfeatures may be implemented using hardware or a combination of hardwareand software. For example, a system like the ones described herein mayinclude various controllers and/or processing devices located at variouspoints. A central computer may coordinate operation among the variouscontrollers or processing devices. The central computer, controllers,and processing devices may execute various software routines to effectcontrol and coordination of testing and calibration.

System operation can be controlled, at least in part, using one or morecomputer program products, e.g., one or more computer program tangiblyembodied in one or more non-transitory machine-readable media, forexecution by, or to control the operation of, one or more dataprocessing apparatus, e.g., a programmable processor, a computer,multiple computers, and/or programmable logic components.

A computer program can be written in any form of programming language,including compiled or interpreted languages, and it can be deployed inany form, including as a stand-alone program or as a module, component,subroutine, or other unit suitable for use in a computing environment. Acomputer program can be deployed to be executed on one computer or onmultiple computers at one site or distributed across multiple sites andinterconnected by a network.

Actions associated with implementing all or part of the operations ofthe particle therapy system described herein can be performed by one ormore programmable processors executing one or more computer programs toperform the functions described herein. All or part of the operationscan be implemented using special purpose logic circuitry, e.g., an FPGA(field programmable gate array) and/or an ASIC (application-specificintegrated circuit).

Processors suitable for the execution of a computer program include, byway of example, both general and special purpose microprocessors, andany one or more processors of any kind of digital computer. Generally, aprocessor will receive instructions and data from a read-only storagearea or a random access storage area or both. Elements of a computer(including a server) include one or more processors for executinginstructions and one or more storage area devices for storinginstructions and data. Generally, a computer will also include, or beoperatively coupled to receive data from, or transfer data to, or both,one or more machine-readable storage media, such as mass PCBs forstoring data, e.g., magnetic, magneto-optical disks, or optical disks.Non-transitory machine-readable storage media suitable for embodyingcomputer program instructions and data include all forms of non-volatilestorage area, including by way of example, semiconductor storage areadevices, e.g., EPROM, EEPROM, and flash storage area devices; magneticdisks, e.g., internal hard disks or removable disks; magneto-opticaldisks; and CD-ROM and DVD-ROM disks.

Any “electrical connection” as used herein may imply a direct physicalconnection or a connection that includes intervening components but thatnevertheless allows electrical signals to flow between connectedcomponents. Any “connection” involving electrical circuitry mentionedherein, unless stated otherwise, is an electrical connection and notnecessarily a direct physical connection regardless of whether the word“electrical” is used to modify “connection”.

Any two more of the foregoing implementations may be used in anappropriate combination in an appropriate particle accelerator (e.g., asynchrocyclotron). Likewise, individual features of any two more of theforegoing implementations may be used in an appropriate combination.

Elements of different implementations described herein may be combinedto form other implementations not specifically set forth above. Elementsmay be left out of the processes, systems, apparatus, etc., describedherein without adversely affecting their operation. Various separateelements may be combined into one or more individual elements to performthe functions described herein.

The example implementations described herein are not limited to use witha particle therapy system or to use with the example particle therapysystems described herein. Rather, the example implementations can beused in any appropriate system that directs accelerated particles to anoutput.

Additional information concerning the design of an exampleimplementation of a particle accelerator that may be used in a system asdescribed herein can be found in U.S. Provisional Application No.60/760,788, entitled “High-Field Superconducting Synchrocyclotron” andfiled Jan. 20, 2006; U.S. patent application Ser. No. 11/463,402,entitled “Magnet Structure For Particle Acceleration” and filed Aug. 9,2006; and U.S. Provisional Application No. 60/850,565, entitled“Cryogenic Vacuum Break Pneumatic Thermal Coupler” and filed Oct. 10,2006, all of which are incorporated herein by reference.

The following applications are incorporated by reference into thesubject application: the U.S. Provisional Application entitled“CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No. 61/707,466),the U.S. Provisional Application entitled “ADJUSTING ENERGY OF APARTICLE BEAM” (Application No. 61/707,515), the U.S. ProvisionalApplication entitled “ADJUSTING COIL POSITION” (Application No.61/707,548), the U.S. Provisional Application entitled “FOCUSING APARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No.61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELDREGENERATOR” (Application No. 61/707,590), the U.S. ProvisionalApplication entitled “FOCUSING A PARTICLE BEAM” (Application No.61/707,704), the U.S. Provisional Application entitled “CONTROLLINGPARTICLE THERAPY (Application No. 61/707,624), and the U.S. ProvisionalApplication entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR”(Application No. 61/707,645).

The following are also incorporated by reference into the subjectapplication: U.S. Pat. No. 7,728,311 which issued on Jun. 1, 2010, U.S.patent application Ser. No. 11/948,359 which was filed on Nov. 30, 2007,U.S. patent application Ser. No. 12/275,103 which was filed on Nov. 20,2008, U.S. patent application Ser. No. 11/948,662 which was filed onNov. 30, 2007, U.S. Provisional Application No. 60/991,454 which wasfiled on Nov. 30, 2007, U.S. Pat. No. 8,003,964 which issued on Aug. 23,2011, U.S. Pat. No. 7,208,748 which issued on Apr. 24, 2007, U.S. Pat.No. 7,402,963 which issued on Jul. 22, 2008, U.S. patent applicationSer. No. 13/148,000 filed Feb. 9, 2010, U.S. patent application Ser. No.11/937,573 filed on Nov. 9, 2007, U.S. patent application Ser. No.11/187,633, titled “A Programmable Radio Frequency Waveform Generatorfor a Synchrocyclotron,” filed Jul. 21, 2005, U.S. ProvisionalApplication No. 60/590,089, filed on Jul. 21, 2004, U.S. patentapplication Ser. No. 10/949,734, titled “A Programmable ParticleScatterer for Radiation Therapy Beam Formation”, filed Sep. 24, 2004,and U.S. Provisional Application No. 60/590,088, filed Jul. 21, 2005.

Any features of the subject application may be combined with one or moreappropriate features of the following: the U.S. Provisional Applicationentitled “CONTROLLING INTENSITY OF A PARTICLE BEAM” (Application No.61/707,466), the U.S. Provisional Application entitled “ADJUSTING ENERGYOF A PARTICLE BEAM” (Application No. 61/707,515), the U.S. ProvisionalApplication entitled “ADJUSTING COIL POSITION” (Application No.61/707,548), the U.S. Provisional Application entitled “FOCUSING APARTICLE BEAM USING MAGNETIC FIELD FLUTTER” (Application No.61/707,572), the U.S. Provisional Application entitled “MAGNETIC FIELDREGENERATOR” (Application No. 61/707,590), the U.S. ProvisionalApplication entitled “FOCUSING A PARTICLE BEAM” (Application No.61/707,704), the U.S. Provisional Application entitled “CONTROLLINGPARTICLE THERAPY (Application No. 61/707,624), and the U.S. ProvisionalApplication entitled “CONTROL SYSTEM FOR A PARTICLE ACCELERATOR”(Application No. 61/707,645), U.S. Pat. No. 7,728,311 which issued onJun. 1, 2010, U.S. patent application Ser. No. 11/948,359 which wasfiled on Nov. 30, 2007, U.S. patent application Ser. No. 12/275,103which was filed on Nov. 20, 2008, U.S. patent application Ser. No.11/948,662 which was filed on Nov. 30, 2007, U.S. ProvisionalApplication No. 60/991,454 which was filed on Nov. 30, 2007, U.S. patentapplication Ser. No. 13/907,601, which was filed on May 31, 2013, U.S.patent application Ser. No. 13/916,401, filed on Jun. 12, 2013, U.S.Pat. No. 8,003,964 which issued on Aug. 23, 2011, U.S. Pat. No.7,208,748 which issued on Apr. 24, 2007, U.S. Pat. No. 7,402,963 whichissued on Jul. 22, 2008, U.S. patent application Ser. No. 13/148,000filed Feb. 9, 2010, U.S. patent application Ser. No. 11/937,573 filed onNov. 9, 2007, U.S. patent application Ser. No. 11/187,633, titled “AProgrammable Radio Frequency Waveform Generator for a Synchrocyclotron,”filed Jul. 21, 2005, U.S. Provisional Application No. 60/590,089, filedon Jul. 21, 2004, U.S. patent application Ser. No. 10/949,734, titled “AProgrammable Particle Scatterer for Radiation Therapy Beam Formation”,filed Sep. 24, 2004, and U.S. Provisional Application No. 60/590,088,filed Jul. 21, 2005.

Other implementations not specifically described herein are also withinthe scope of the following claims.

What is claimed is:
 1. A particle therapy system comprising: a particleaccelerator to output a beam of charged particles, the particleaccelerator being configured to produce the beam in accordance with aradio frequency (RF) cycle; and a scanning system to perform scanning onat least part of an irradiation target, the scanning system comprising:a scanning magnet to move the beam across the at least part of theirradiation target during scanning, where movement of the beam is basedon a current of the scanning magnet; and a control system (i) to controlthe current of the scanning magnet independently of the RF cycle inorder to move the beam across the at least part of the irradiationtarget to deliver doses of charged particles, (ii) for positions on theirradiation target at which the particle beam delivers a dose of chargedparticles, to store information identifying a location and an amount ofthe dose delivered, (iii) to compare a cumulative dose delivered at eachposition to a target cumulative dose, and (iv) if the cumulative dosedoes not match the target cumulative dose at specific positions, tocontrol the current of the scanning magnet in order to move the beam soas to deliver additional dose to the specific positions.
 2. The particletherapy system of claim 1, wherein the control system is configured tomeasure the cumulative dose delivered at each position; and whereinmeasuring the cumulative dose delivered at each position issubstantially synchronous with the RF cycle.
 3. The particle therapysystem of claim 1, wherein the control system is configured to measurethe cumulative dose delivered at each position; and wherein measuringthe cumulative dose delivered at each position is substantiallysynchronous with delivery of dose at each position.
 4. The particletherapy system of claim 1, wherein the information comprises an amountof dose delivered at each position and at least one of: a location ofeach position within the irradiation target or a magnet currentcorresponding to each position within the irradiation target.
 5. Theparticle therapy system of claim 1, wherein the location corresponds tothree-dimensional coordinates for the irradiation target.
 6. Theparticle therapy system of claim 1, wherein the particle therapy systemfurther comprises: memory to store a treatment plan that identifies, foreach position, a target cumulative dose of charged particles, thetreatment plan omitting information about individual doses of chargedparticles to be delivered to individual positions during scanning toachieve the target cumulative dose.
 7. The particle therapy system ofclaim 1, wherein the scanning system further comprises: a degrader tochange an energy of the beam prior to the beam reaching the irradiationtarget, the degrader being down-beam of the scanning magnet relative tothe particle accelerator; wherein the control system is configured tocontrol movement of at least part of the degrader into, or out of, apath of the beam during scanning in order to affect the energy of thebeam and thereby set a layer of the irradiation target to which chargedparticles are to be delivered.
 8. The particle therapy system of claim7, wherein the particle accelerator comprises an ion source to providepulses of charged particles, the pulses of charged particles forming thebeam; and wherein, during at least part of the movement of the degrader,the ion source is deactivated.
 9. The particle therapy system of claim7, wherein the particle accelerator comprises: an ion source to providepulses of charged particles, the pulses of charged particles forming thebeam; and a voltage source to provide an RF voltage at the RF cycle to acavity to accelerate the charged particles, the cavity having a magneticfield for causing the charged particles to move orbitally within thecavity; wherein, during at least part of the movement of the degrader,the voltage source is deactivated.
 10. The particle therapy system ofclaim 9, wherein, during the movement of the at least part of thedegrader, the ion source is deactivated at a same time that the voltagesource is deactivated.
 11. The particle therapy system of claim 1,wherein the particle accelerator is a variable-energy particleaccelerator; and wherein the control system is configured to set anenergy level of the particle accelerator prior to scanning, the energylevel comprising an energy level of the beam.
 12. The particle therapysystem of claim 1, wherein the particle accelerator is a variable-energyparticle accelerator; and wherein the control system is configured tochange an energy level of the particle accelerator during scanning, theenergy level comprising an energy level of the beam.
 13. The particletherapy system of claim 1, wherein, for a target position at which theparticle beam delivers a dose of charged particles, multiple individualdeliveries of the dose of charged particles comprise a percentage of atotal cumulative dose at the target position.
 14. The particle therapysystem of claim 13, wherein the percentage is less than 100% of thetotal cumulative dose at the target position.
 15. The particle therapysystem of claim 13, wherein the percentage is about 100% of the totalcumulative dose at the target position.
 16. The particle therapy systemof claim 1, wherein the scanning magnet has an air core.
 17. Theparticle therapy system of claim 1, wherein the scanning magnet has aferromagnetic core.
 18. The particle therapy system of claim 1, furthercomprising: a gantry on which the particle accelerator is mounted, thegantry for rotating the particle accelerator around the irradiationtarget; wherein the control system is configured to control the scanningsystem based on a rotational angle of the gantry.
 19. The particletherapy system of claim 1, wherein the particle accelerator is asynchrocyclotron; and wherein the beam is a pulsed beam comprised ofpulses of charged particles, the charged particles in the beam beingoutput from the synchrocyclotron in accordance with the RF cycle.
 20. Aparticle therapy system comprising: a particle accelerator to output abeam of charged particles, the particle accelerator being configured toproduce the beam in accordance with a radio frequency (RF) cycle; and ascanning system to perform scanning on least part of an irradiationtarget, the scanning system comprising: a scanning magnet to move thebeam across the at least part of the irradiation target during scanning,where movement of the beam is based on a current of the scanning magnet;and a control system (i) to control the current of the scanning magnetindependently of the RF cycle in order to move the beam across the atleast part of the irradiation target so as to deliver doses of chargedparticles to positions on the irradiation target; and (ii) to determine,in synchronism with delivery of a dose of charged particles to aposition on the irradiation target, information identifying an amount ofthe dose actually delivered to the position.
 21. The particle therapysystem of claim 20, wherein the particle accelerator is configured tooutput pulses of charged particles in accordance with the RF cycle, thepulses of charged particles forming the beam.
 22. The particle therapysystem of claim 20, further comprising: a gantry on which the particleaccelerator is mounted, the gantry for rotating the particle acceleratoraround the irradiation target; wherein the control system is configuredto control the scanning system based on a rotational angle of thegantry.
 23. The particle therapy system of claim 22, wherein thescanning magnet comprises an air core; and wherein controlling thescanning system based on a rotational angle of the gantry comprisescontrolling the current of the scanning magnet based on the rotationalangle of the gantry.
 24. The particle therapy system of claim 20,further comprising: a gantry on which the particle accelerator ismounted, the gantry for rotating the particle accelerator around theirradiation target; wherein the scanning magnet comprises aferromagnetic core.
 25. The particle therapy system of claim 20, whereinthere is a linear correlation between the current of the scanning magnetand the movement of the beam during at least part of the movement of thebeam.
 26. The particle therapy system of claim 20, wherein the controlsystem is configured to measure a cumulative dose of charged particlesdelivered to positions on the irradiation target; and wherein measuringthe cumulative dose delivered at each of the positions is substantiallysynchronous with the RF cycle.
 27. The particle therapy system of claim20, wherein the scanning magnet comprises a core comprised of both airand ferromagnetic material.
 28. The particle therapy system of claim 27,further comprising: a gantry on which the particle accelerator ismounted, the gantry for rotating the particle accelerator around theirradiation target; wherein the control system is configured to controlthe scanning system based on a rotational angle of the gantry.
 29. Theparticle therapy system of claim 22, wherein controlling the scanningsystem based on a rotational angle of the gantry comprises controllingthe current of the scanning magnet based on the rotational angle of thegantry.
 30. The particle therapy system of claim 20, further comprising:a range modulator comprising plates configured to move into, and out of,a path of the particle beam during scanning of the irradiation target.31. The particle therapy system of claim 20, wherein the control systemis configured to change an intensity of the beam for different scans ofthe particle beam.
 32. The particle therapy system of claim 31, whereinthe control system is configured to cause the beam to move across the atleast part of the irradiation target multiple times; and wherein, fordifferent times that the beam moves across the at least part of theirradiation target, the control system is configured to change theintensity of the beam.
 33. The particle therapy system of claim 20,wherein the particle therapy system further comprises: memory to store atreatment plan that identifies, for each position to which a dose is tobe delivered on the irradiation target, a target cumulative dose ofcharged particles, the treatment plan omitting information aboutindividual doses of charged particles to be delivered to individualpositions during scanning to achieve the target cumulative dose; whereinthe scanning system is configured to perform scanning based on thetreatment plan.
 34. The particle therapy system of claim 20, whereindetermining the information identifying an amount of the dose actuallydelivered to the position comprises measuring the amount of the doseactually delivered to the position as the dose is delivered.